Guided Mode Resonant Filter Biosensor Using a Linear Grating Surface Structure

ABSTRACT

Methods and compositions are provided for detecting biomolecular interactions. The use of labels is not required and the methods can be performed in a high-throughput manner. The invention also provides optical devices useful as narrow band filters.

PRIORITY

This application is a divisional application of U.S. Ser. No.11/201,237, filed Aug. 10, 2005 (allowed), which is acontinuation-in-part of U.S. application Ser. No. 10/059,060, filed Jan.28, 2002, now U.S. Pat. No. 7,070,987. This application is also acontinuation-in-part of U.S. application Ser. No. 09/930,352, filed Aug.15, 2001, now U.S. Pat. No. 7,094,595, which claims the benefit of U.S.provisional application 60/244,312 filed Oct. 30, 2000; U.S. provisionalapplication 60/283,314 filed Apr. 12, 2001; and U.S. provisionalapplication 60/303,028 filed Jul. 3, 2001.

TECHNICAL AREA OF THE INVENTION

The invention relates to compositions and methods for detectingbiomolecular interactions. The detection can occur without the use oflabels and can be done in a high-throughput manner. The invention alsorelates to optical devices.

BACKGROUND OF THE INVENTION

With the completion of the sequencing of the human genome, one of thenext grand challenges of molecular biology will be to understand how themany protein targets encoded by DNA interact with other proteins, smallmolecule pharmaceutical candidates, and a large host of enzymes andinhibitors. See e.g., Pandey & Mann, “Proteomics to study genes andgenomes,” Nature, 405, p. 837-846, 2000; Leigh Anderson et al.,“Proteomics: applications in basic and applied biology,” Current Opinionin Biotechnology, 11, p. 408-412, 2000; Patterson, “Proteomics: theindustrialization of protein chemistry,” Current Opinion inBiotechnology, 11, p. 413-418, 2000; MacBeath & Schreiber, “PrintingProteins as Microarrays for High-Throughput Function Determination,”Science, 289, p. 1760-1763, 2000; De Wildt et al., “Antibody arrays forhigh-throughput screening of antibody-antigen interactions,” NatureBiotechnology, 18, p. 989-994, 2000. To this end, tools that have theability to simultaneously quantify many different biomolecularinteractions with high sensitivity will find application inpharmaceutical discovery, proteomics, and diagnostics. Further, forthese tools to find widespread use, they must be simple to use,inexpensive to own and operate, and applicable to a wide range ofanalytes that can include, for example, polynucleotides, peptides, smallproteins, antibodies, and even entire cells.

Biosensors have been developed to detect a variety of biomolecularcomplexes including oligonucleotides, antibody-antigen interactions,hormone-receptor interactions, and enzyme-substrate interactions. Ingeneral, biosensors consist of two components: a highly specificrecognition element and a transducer that converts the molecularrecognition event into a quantifiable signal. Signal transduction hasbeen accomplished by many methods, including fluorescence,interferometry (Jenison et al, “Interference-based detection of nucleicacid targets on optically coated silicon,” Nature Biotechnology, 19, p.62-65; Lin et al, “A porous silicon-based optical interferometricbiosensor,” Science, 278, p. 840-843, 1997), and gravimetry (A.Cunningham, Bioanalytical Sensors, John Wiley & Sons (1998)).

Of the optically-based transduction methods, direct methods that do notrequire labeling of analytes with fluorescent compounds are of interestdue to the relative assay simplicity and ability to study theinteraction of small molecules and proteins that are not readilylabeled. Direct optical methods include surface plasmon resonance (SPR)(Jordan & Corn, “Surface Plasmon Resonance Imaging Measurements ofElectrostatic Biopolymer Adsorption onto Chemically Modified GoldSurfaces,” Anal. Chem., 69:1449-1456 (1997), (grating couplers (Morhardet al., “Immobilization of antibodies in micropatterns for celldetection by optical diffraction,” Sensors and Actuators B, 70, p.232-242, 2000), ellipsometry (Jin et al., “A biosensor concept based onimaging ellipsometry for visualization of biomolecular interactions,”Analytical Biochemistry, 232, p. 69-72, 1995), evanascent wave devices(Huber et al., “Direct optical immunosensing (sensitivity andselectivity),” Sensors and Actuators B, 6, p. 122-126, 1992), andreflectometry (Brecht & Gauglitz, “Optical probes and transducers,”Biosensors and Bioelectronics, 10, p. 923-936, 1995). Theoreticallypredicted detection limits of these detection methods have beendetermined and experimentally confirmed to be feasible down todiagnostically relevant concentration ranges. However, to date, thesemethods have yet to yield commercially available high-throughputinstruments that can perform high sensitivity assays without any type oflabel in a format that is readily compatible with the microtiterplate-based or microarray-based infrastructure that is most often usedfor high-throughput biomolecular interaction analysis. Therefore, thereis a need in the art for compositions and methods that can achieve thesegoals.

SUMMARY OF THE INVENTION

It is an object of the invention to provide compositions and methods fordetecting binding of one or more specific binding substances to theirrespective binding partners. This and other objects of the invention areprovided by one or more of the embodiments described below.

One embodiment of the invention provides a biosensor. The biosensorcomprises a one-dimensional grating layer comprised of a material havinga high refractive index, a low refractive index material layer thatsupports the one-dimensional grating layer, and one or more specificbinding substances immobilized on the surface of the one-dimensionalgrating layer opposite of the low refractive index material layer. Whenthe biosensor is illuminated a resonant grating effect is produced on areflected radiation spectrum. The cross-sectional period of theone-dimensional grating is less than the wavelength of the resonantgrating effect. In another embodiment, the biosensor comprises aone-dimensional grating surface structure comprised of a material havinga low refractive index, a high refractive index material layer that isapplied on top of the low refractive index one-dimensional gratinglayer, and one or more specific binding substances immobilized on asurface of the high refractive index layer opposite of theone-dimensional grating surface structure comprised of a material havinga low refractive index.

The cross-sectional profile of the one-dimensional grating can betriangular, sinusoidal, trapezoidal, rectangular, stepped, v-shaped,u-shaped, upside-down v-shaped, upside-down u-shaped, or square. Anarrow band of optical wavelengths is reflected from the biosensor whenthe biosensor is illuminated with a broad band of optical wavelengths.

The low refractive index material of the biosensor can comprise glass,plastic, polymer, or epoxy. The high refractive index material can beselected from the group consisting of zinc sulfide, titanium dioxide,indium tin oxide, tantalum oxide, and silicon nitride. Theone-dimensional grating can have a period of about 0.01 microns to about1 micron and a depth of about 0.01 microns to about 1 micron.

The one or more specific binding substances can be arranged in an arrayof distinct locations. The distinct locations can define a microarrayspot of about 50-500 microns in diameter. The one or more specificbinding substances can be immobilized on the high refractive indexmaterial by physical adsorption or by chemical binding. The one or morespecific binding substances can be bound to their binding partners. Theone or more specific binding substances or binding partners can beselected from the group consisting of nucleic acids, polypeptides,antigens, polyclonal antibodies, monoclonal antibodies, single chainantibodies (scFv), F(ab) fragments, F(ab′)₂ fragments, Fv fragments,small organic molecules, cells, viruses, bacteria, polymers, proteinsolutions, peptide solutions, single- or double-stranded DNA solutions,RNA solutions, solutions containing compounds from a combinatorialchemical library and biological samples. The biological sample can beselected from the group consisting of, blood, plasma, serum,gastrointestinal secretions, homogenates of tissues or tumors, synovialfluid, feces, saliva, sputum, cyst fluid, amniotic fluid, cerebrospinalfluid, peritoneal fluid, lung lavage fluid, semen, lymphatic fluid,tears, and prostatitc fluid.

A biosensor can comprise the internal surface of a liquid-containingvessel, such as a microtiter plate, a test tube, a petri dish and amicrofluidic channel.

Another embodiment of the invention provides a detection systemcomprising a biosensor, a light source that directs light to thebiosensor, and a detector that detects light reflected from thebiosensor, wherein a polarizing filter occurs between the light sourceand the biosensor.

Yet another embodiment of the invention provides a method of detectingthe binding of one or more specific binding substances to theirrespective binding partners. The method comprises applying one or morebinding partners to a biosensor, illuminating the biosensor with light;and detecting a peak wavelength value (PWV). If the one or more specificbinding substances have bound to their respective binding partners, thenthe PWV is shifted.

Still another embodiment of the invention provides a method of detectingthe binding of one or more specific binding substances to theirrespective binding partners. The method comprises applying one or morebinding partners to a biosensor, wherein the high refractive indexmaterial is coated with an array of distinct locations containing theone or more specific binding substances, illuminating each distinctlocation of the biosensor with light; and detecting peak wavelengthvalue (PWV) for each distinct location of the biosensor. If the one ormore specific binding substances have bound to their respective bindingpartners at a distinct location, then the PWV is shifted.

Even another embodiment of the invention provides a method of detectingactivity of an enzyme. The method comprises applying one or more enzymesto a biosensor, washing the biosensor, illuminating the biosensor withlight, and detecting a PWV. If the one or more enzymes have altered theone or more specific binding substances of the biosensor by enzymaticactivity, then the PWV is shifted.

Another embodiment provides a method of measuring the amount of one ormore binding partners in a test sample. The method comprisesilluminating a biosensor with light, detecting a PWV from the biosensor,applying a test sample comprising one or more binding partners to thebiosensor, illuminating the biosensor with light; and, detecting a PWVfrom the biosensor. The difference in PWVs is a measurement of theamount of one or more binding partners in the test sample.

Still another embodiment of the invention provides a method of detectingthe binding of one or more specific binding substances to theirrespective binding partners. The method comprises applying one or morebinding partners comprising one or more tags to a biosensor,illuminating the biosensor with light; and detecting a PWV from thebiosensor. If the one or more specific binding substances have bound totheir respective binding partners, then the reflected wavelength oflight is shifted. The one or more tags can be selected from the groupconsisting of biotin,succinimidyl-6-[a-methyl-a-(2-pyridyl-dithio)toluamido]hexanoate (SMPT),dimethylpimelimidate (DMP), and histidine. The one or more tags can bereacted with a composition selected from the group consisting ofstreptavidin, horseradish peroxidase, and streptavidin coatednanoparticles, before the step of illuminating the biosensor with light.

Another embodiment of the invention provides a biosensor comprising aone-dimensional or two-dimensional grating layer comprised of a materialhaving a high refractive index, a low refractive index material layerthat supports the one-dimensional or two-dimensional grating layer; asurface modification layer on a surface of the one-dimensional ortwo-dimensional grating layer opposite of the low refractive indexmaterial layer; and one or more specific binding substances immobilizedon a surface of the surface modification layer opposite of theone-dimensional or two-dimensional grating layer. When the biosensor isilluminated a resonant grating effect is produced on a reflectedradiation spectrum. The surface modification layer can be comprised ofsilicon oxide. The thickness of the surface modification layer can beabout 5 nm to about 15 nm.

Another embodiment of the invention provides a biosensor comprising agrating layer comprising a one-dimensional or two-dimensional grating ona first surface; a interfacial layer on the first surface of the gratinglayer, a high refractive index material layer on the surface of theinterfacial layer opposite of the grating layer, and one or morespecific binding substances immobilized on a surface of the highrefractive index material layer opposite of the interfacial layer. Whenthe biosensor is illuminated a resonant grating effect is produced on areflected radiation spectrum. The interfacial layer can be comprised ofa material selected from the group consisting of silicon oxide, siliconoxynitride, borosilicate glass, phosophosilicate glass, pyrex, glass,and a metal oxide. The interfacial layer can be about 1 nm to about 200nm thick.

Therefore, unlike surface plasmon resonance, resonant mirrors, andwaveguide biosensors, the described compositions and methods enable manythousands of individual binding reactions to take place simultaneouslyupon the biosensor surface. This technology is useful in applicationswhere large numbers of biomolecular interactions are measured inparallel, particularly when molecular labels will alter or inhibit thefunctionality of the molecules under study. High-throughput screening ofpharmaceutical compound libraries with protein targets, and microarrayscreening of protein-protein interactions for proteomics are examples ofapplications that require the sensitivity and throughput afforded bythis approach. A biosensor of the invention can be manufactured, forexample, in large areas using a plastic embossing process, or an epoxyreplication process, and thus can be inexpensively incorporated intocommon disposable laboratory assay platforms such as microtiter platesand microarray slides.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1A shows a schematic diagram of one embodiment of an opticalgrating structure used for a calorimetric resonant reflectancebiosensor. n_(substrate) represents substrate material. n₁ representsthe refractive index of a cover layer. n₂ represents the refractiveindex of a one- or two-dimensional grating. n_(bio) represents therefractive index of one or more specific binding substances. t₁represents the thickness of the cover layer. t₂ represents the thicknessof the grating. t_(bio) represents the thickness of the layer of one ormore specific binding substances. FIG. 1 shows a cross-sectional view ofa biosensor.

FIG. 2 shows a schematic drawing of a one-dimensional linear gratingsurface structure.

FIG. 3A-B shows a two-dimensional grating comprising a rectangular gridof squares (FIG. 3A) or holes (FIG. 3B).

FIG. 4 shows a biosensor cross-section profile utilizing a sinusoidallyvarying grating profile.

FIG. 5 shows a biosensor cross-section profile in which an embossedsubstrate is coated with a higher refractive index material such as ZnSor SiN. An optional cover layer of low refractive index material, forexample, epoxy or SOG is layered on top of the higher refractive indexmaterial and one or more specific binding substances are immobilized onthe cover layer.

FIG. 6 shows three types of surface activation chemistry (Amine,Aldehyde, and Nickel) with corresponding chemical linker molecules thatcan be used to covalently attach various types of biomolecule receptorsto a biosensor.

FIG. 7A-C shows methods that can be used to amplify the mass of abinding partner such as detected DNA or detected protein on the surfaceof a biosensor.

FIG. 8 shows a graphic representation of how adsorbed material, such asa protein monolayer, will increase the reflected wavelength of on a SRVDbiosensor.

FIG. 9 shows an example of a biosensor used as a microarray.

FIG. 10A-B shows two biosensor formats that can incorporate acalorimetric resonant reflectance biosensor. FIG. 10A shows a biosensorthat is incorporated into a microtitre plate.

FIG. 10B shows a biosensor in a microarray slide format.

FIG. 11 shows an array of arrays concept for using a biosensor platformto perform assays with higher density and throughput.

FIG. 12 shows a diagram of an array of biosensor electrodes. A singleelectrode can comprise a region that contains many grating periods andseveral separate grating regions can occur on the same substratesurface.

FIG. 13 shows a SEM photograph showing the separate grating regions ofan array of biosensor electrodes.

FIG. 14 shows a biosensor upper surface immersed in a liquid sample. Anelectrical potential can be applied to the biosensor that is capable ofattracting or repelling a biomolecule near the electrode surface.

FIG. 15 shows a biosensor upper surface immersed in a liquid sample. Apositive voltage is applied to an electrode and the electronegativebiomolecules are attracted to the biosensor surface.

FIG. 16 shows a biosensor upper surface immersed in a liquid sample. Anegative voltage is applied to an electrode and the electronegativebiomolecules are repelled from the biosensor surface using a negativeelectrode voltage.

FIG. 17 demonstrates an example of a biosensor that occurs on the tip ofa fiber probe for in vivo detection of biochemical substances.

FIG. 18 shows an example of the use of two coupled fibers to illuminateand collect reflected light from a biosensor.

FIG. 19 shows resonance wavelength of a biosensor as a function ofincident angle of detection beam.

FIG. 20 shows an example of the use of a beam splitter to enableilluminating and reflected light to share a common collimated opticalpath to a biosensor.

FIG. 21 shows an example of a system for angular scanning of abiosensor.

FIG. 22 shows SEM photographs of a photoresist grating structure in planview (center and upper right) and cross-section (lower right).

FIG. 23 shows a SEM cross-section photograph of a grating structureafter spin-on glass is applied over a silicon nitride grating.

FIG. 24 shows examples of biosensor chips (1.5×10.5-inch). Circularareas are regions where the resonant structure is defined.

FIG. 25 shows response as a function of wavelength of a biosensor thatBSA had been deposited at high concentration, measured in air. Beforeprotein deposition, the resonant wavelength of the biosensor is 380 nmand is not observable with the instrument used for this experiment.

FIG. 26 shows response as a function of wavelength comparing anuntreated biosensor with one upon which BSA had been deposited. Bothmeasurements were taken with water on the biosensor's surface.

FIG. 27 shows response as a function of wavelength of a biosensor thatBorrelia bacteria has been deposited at high concentration and measuredin water.

FIG. 28 shows a computer simulation of a biosensor demonstrating theshift of resonance to longer wavelengths as biomolecules are depositedon the surface.

FIG. 29 shows a computer simulation demonstrating the dependence of peakreflected wavelength on protein coating thickness. This particularbiosensor has a dynamic range of 250 nm deposited biomaterial before theresponse begins to saturate.

FIG. 30 shows an embodiment of a biosensor. n_(substrate) represents therefractive index of a substrate. n₁ represents the refractive index ofan optional optical cover layer. n₂ represents the refractive index of aone- or two-dimensional grating. n₃ represents the refractive index of ahigh refractive index material such as silicon nitride. n_(bio)represents the refractive index of one or more specific bindingsubstances. t₁ represents the thickness of a cover layer. t₂ representsthe thickness of a one- or two-dimensional grating. t₃ represents thethickness of a high refractive index material. t_(bio) represents thethickness of a specific binding substance layer.

FIG. 31 shows reflected intensity as a function of wavelength for aresonant grating structure when various thicknesses of protein areincorporated onto the upper surface.

FIG. 32 shows a linear relationship between reflected wavelength andprotein coating thickness for a biosensor shown in FIG. 30.

FIG. 33 shows instrumentation that can be used to read output of abiosensor. A collimated light source is directed at a biosensor surfaceat normal incidence through an optical fiber, while a second parallelfiber collects the light reflected at normal incidence. A spectrometerrecords the reflectance as a function of wavelength.

FIG. 34 shows the measured reflectance spectra of a biosensor.

FIG. 35 shows dependence of peak resonant wavelength measured in liquidupon the concentration of protein BSA dissolved in water.

FIG. 36 shows dependence of peak resonance wavelength on theconcentration of BSA dissolved in PBS, which was then allowed to dry ona biosensor surface.

FIG. 37A-B. FIG. 37A shows a measurement of peak resonant wavelengthshift caused by attachment of a streptavidin receptor layer andsubsequent detection of a biotinylated IgG. FIG. 37B shows a schematicdemonstration of molecules bound to a biosensor.

FIG. 38A-B. FIG. 38A shows results of streptavidin detection at variousconcentrations for a biosensor that has been activated with NH₂ surfacechemistry linked to a biotin receptor molecule. FIG. 38B shows aschematic demonstration of molecules bound to a biosensor.

FIG. 39A-B. FIG. 39A shows an assay for detection of anti-goat IgG usinga goat antibody receptor molecule. BSA blocking of a detection surfaceyields a clearly measurable background signal due to the mass of BSAincorporated on the biosensor. A 66 nM concentration of anti-goat IgG iseasily measured above the background signal. FIG. 39B shows a schematicdemonstration of molecules bound to a biosensor.

FIG. 40A-B. FIG. 40A shows a nonlabeled ELISA assay for interferon-gamma(INF-gamma) using an anti-human IgG INF-gamma receptor molecule, and aneural growth factor (NGF) negative control. FIG. 40B shows a schematicdemonstration of molecules bound to a biosensor.

FIG. 41A-B. FIG. 41A shows detection of a 5-amino acid peptide (MW=860)and subsequent cleavage of a pNA label (MW=130) using enzyme caspase-3.FIG. 41B shows a schematic demonstration of molecules bound to abiosensor.

FIG. 42A-B. FIG. 42A shows resonant peak in liquid during continuousmonitoring of the binding of three separate protein layers. FIG. 42Bshows a schematic demonstration of molecules bound to a biosensor.

FIG. 43A-B. FIG. 43A shows endpoint resonant frequencies mathematicallydetermined from the data shown in FIG. 42. FIG. 43B shows a schematicdemonstration of molecules bound to a biosensor.

FIG. 44A-B. FIG. 44A shows kinetic binding measurement of IgG binding.FIG. 44B shows a schematic demonstration of molecules bound to abiosensor.

FIG. 45A-B. FIG. 45A shows kinetic measurement of a protease thatcleaves bound protein from a biosensor surface. FIG. 45B shows aschematic demonstration of molecules bound to a biosensor.

FIG. 46 shows comparison of mathematical fit of parabolic andexponential functions to spectrometer data from a resonant peak. Theexponential curve fit is used to mathematically determine a peakresonant wavelength.

FIG. 47 shows sensitivity of the mathematically determined peak resonantwavelength to artificially added noise in the measured spectrum.

FIG. 48 shows a resonant optical biosensor incorporating an electricallyconducting material.

FIG. 49 shows a resonant reflection or transmission filter structureconsisting of a set of concentric rings.

FIG. 50 shows a resonant reflective or transmission filter structurecomprising a hexagonal grid of holes (or a hexagonal grid of posts) thatclosely approximates the concentric circle structure of FIG. 49 withoutrequiring the illumination beam to be centered upon any particularlocation of the grid.

FIG. 51 shows a plot of the peak resonant wavelength values for testsolutions. The avidin solution was taken as the baseline reference forcomparison to the Avidin+BSA and Avidin+b-BSA solutions. Addition of BSAto avidin results in only a small resonant wavelength increase, as thetwo proteins are not expected to interact. However, because biotin andavidin bind strongly (Kd=10⁻¹⁵M), the avidin+b-BSA solution will containlarger bound protein complexes. The peak resonant wavelength value ofthe avidin+b-BSA solution thus provides a large shift compared toavidin+BSA.

FIG. 52 shows a schematic diagram of a detection system.

FIG. 53A-B shows a fabrication process used to produce the biosensor andcross-section of a one-dimensional linear grating sensor. FIG. 53A showsa silicon master wafer used to replicate the biosensor structure into athin film of epoxy between the silicon and a sheet of plastic film.After the epoxy is cured, the plastic sheet is peeled away. To completesensor fabrication (FIG. 53B), a thin film of high refractive indexdielectric material such as silicon nitride, titanium oxide, tantalumoxide, or zinc sulfide is deposited over the structure.

FIG. 54A-C shows a linear grating structure (FIG. 54A; top view) used toproduce the one-dimensional linear grating guided mode resonant filter“master” structure. First, an 8-inch diameter silicon “master” wafer isproduced. The 550 nm period linear grating structure is defined inphotoresist using deep-UV photolithography by stepping and repeating theexposure of a 9 mm diameter circular grating reticle over the surface ofa photoresist-coated silicon wafer, as shown in FIG. 54B. FIG. 54C showsthat the exposure step/repeat procedure produced patterns for twostandard format 96-well microtiter plates with 8 rows and 12 columnseach. The exposed photoresist was developed, and the grating structurewas permanently transferred to the silicon wafer using a reactive ionetch with a depth of ˜200 nm. After etching, the photoresist wasremoved.

FIG. 55 shows instrumentation used to illuminate and read output of abiosensor structure. The probe head contains two optical fibers. Thefirst fiber is connected to a white light source to cast a small spot ofpolarized collimated light on the biosensor surface. The second fibercollects reflected light for analysis by a spectrometer.

FIG. 56 shows reflected intensity as a function of wavelength for aone-dimensional linear grating surface biosensor structure within amicrotiter plate well filled with water.

FIG. 57 demonstrates peak wavelength shift relative to a cleanone-dimensional linear grating surface biosensor structure for threebiosensor surface activation states. The error bars indicate thestandard deviation of the shift over seven separate sensor wells.

FIG. 58A-C shows the exposure of NH₂, PEG, and PEG-Biotin activatedone-dimensional linear grating surface biosensor structures to sevenconcentrations of anti-biotin IgG. The NH₂ surface (FIG. 58A) displayslow levels of nonspecific protein binding at high protein exposureconcentrations, while the PEG surface (FIG. 58B) displays low levels ofnonspecific binding. The PEG-Biotin (FIG. 58C) surface has a strongbinding interaction with the anti-biotin IgG.

FIG. 59 shows peak wavelength value shift as a function of anti-biotinIgG concentration for PEG-Biotin activated wells after a 20-minuteincubation. The plotted line indicates a least-squared fit linearfunction.

FIG. 60 demonstrates the effect of a surface modification layer onspecific binding substance immobilization onto the surface of abiosensor.

FIG. 61 shows water stability test results for biosensors with andwithout an interfacial layer. The addition of an interfacial layersignificantly improved stability of a biosensor in aqueous solutions.

DETAILED DESCRIPTION OF THE INVENTION

Subwavelength Structured Surface (SWS) Biosensor

In one embodiment of the invention, a subwavelength structured surface(SWS) is used to create a sharp optical resonant reflection at aparticular wavelength that can be used to track with high sensitivitythe interaction of biological materials, such as specific bindingsubstances or binding partners or both. A colormetric resonantdiffractive grating surface acts as a surface binding platform forspecific binding substances.

SWSs are an unconventional type of diffractive optic that can mimic theeffect of thin-film coatings. (Peng & Morris, J. Opt. Soc. Am. A, Vol.13, No. 5, p. 993, May 1996; Magnusson, & Wang, Appl. Phys. Lett., 61,No. 9, p. 1022, August, 1992; Peng & Morris, Optics Letters, Vol. 21,No. 8, p. 549, April, 1996). A SWS structure comprises a surface-reliefgrating, such as a one-dimensional, two-dimensional, or threedimensional grating in which the grating period is small compared to thewavelength of incident light.

The reflected or transmitted color of this structure can be modulated bythe addition of molecules such as specific binding substances, bindingpartners, or both, or inorganic molecules to the upper surface of thecover layer or the grating surface. The dielectric susceptibility of theadded molecules results in a modification of the wavelength at whichmaximum reflectance or transmittance will occur.

In one embodiment, a biosensor, when illuminated with white light, isdesigned to reflect only a single wavelength. When specific bindingsubstances are attached to the surface of the biosensor, the reflectedwavelength (color) is shifted due to the change of the optical path oflight that is coupled into the grating. By linking specific bindingsubstances to a biosensor surface, complementary binding partnermolecules can be detected without the use of any kind of fluorescentprobe or particle label. The detection technique is capable of resolvingchanges of, for example, ˜0.1 nm thickness of protein binding, and canbe performed with the biosensor surface either immersed in fluid ordried.

A detection system consists of, for example, a light source thatilluminates a small spot of a biosensor at normal incidence through, forexample, a fiber optic probe, and a spectrometer that collects thereflected light through, for example, a second fiber optic probe also atnormal incidence. Because no physical contact occurs between theexcitation/detection system and the biosensor surface, no specialcoupling prisms are required and the biosensor can be easily adapted toany commonly used assay platform including, for example, microtiterplates and microarray slides. A single spectrometer reading can beperformed in several milliseconds, thus it is possible to quicklymeasure a large number of molecular interactions taking place inparallel upon a biosensor surface, and to monitor reaction kinetics inreal time.

This technology is useful in applications where large numbers ofbiomolecular interactions are measured in parallel, particularly whenmolecular labels would alter or inhibit the functionality of themolecules under study. High-throughput screening of pharmaceuticalcompound libraries with protein targets, and microarray screening ofprotein-protein interactions for proteomics are examples of applicationsthat require the sensitivity and throughput afforded by the compositionsand methods of the invention.

A schematic diagram of an example of a SWS structure is shown in FIG. 1.In FIG. 1, n_(substrate) represents a substrate material. n₁ representsthe refractive index of an optional cover layer. n₂ represents therefractive index of a two-dimensional grating. N_(bio) represents therefractive index of one or more specific binding substances. t₁represents the thickness of the cover layer above the two-dimensionalgrating structure. t₂ represents the thickness of the grating. t_(bio)represents the thickness of the layer of one or more specific bindingsubstances. In one embodiment, are n2>n1. (see FIG. 1). Layerthicknesses (i.e. cover layer, one or more specific binding substances,or a grating) are selected to achieve resonant wavelength sensitivity toadditional molecules on the top surface The grating period is selectedto achieve resonance at a desired wavelength.

One embodiment of the invention provides a SWS biosensor. A SWSbiosensor comprises a one-dimensional or two-dimensional grating, asubstrate layer that supports the grating, and one or more specificbinding substances immobilized on the surface of the grating opposite ofthe substrate layer.

A one-dimensional or two-dimensional grating can be comprised of amaterial, including, for example, zinc sulfide, titanium dioxide,tantalum oxide, and silicon nitride. A cross-sectional profile of thegrating can comprise any periodically repeating function, for example, a“square-wave.” A grating can be comprised of a repeating pattern ofshapes selected from the group consisting of continuous parallel linessquares, circles, ellipses, triangles, trapezoids, sinusoidal waves,ovals, rectangles, and hexagons. A sinusoidal cross-sectional profile ispreferable for manufacturing applications that require embossing of agrating shape into a soft material such as plastic, or replicating agrating surface into a material such as epoxy. In one embodiment of theinvention, the depth of the grating is about 0.01 micron to about 1micron and the period of the grating is about 0.01 micron to about 1micron.

A SWS biosensor can also comprise a one-dimensional linear gratingsurface structure, i.e., a series of parallel lines or grooves. Seee.g., FIG. 54. A one-dimensional linear grating is sufficient forproducing the guided mode resonant filter effect. While atwo-dimensional grating has features in two lateral directions acrossthe plane of the sensor surface that are both subwavelength, thecross-section of a one-dimensional grating is only subwavelength in onelateral direction, while the long dimension can be greater thanwavelength of the resonant grating effect. A one-dimensional gratingbiosensor can comprise a high refractive index material which is coatedas a thin film over a layer of lower refractive index material with thesurface structure of a one-dimensional grating. See FIG. 53.Alternatively, a one dimensional grating biosensor can comprise a lowrefractive index material substrate, upon which a high refractive indexthin film material has been patterned into the surface structure of aone-dimensional grating. The low refractive index material can be glass,plastic, polymer, or cured epoxy. The high refractive index materialmust have a refractive index that is greater than the low refractiveindex material. The high refractive index material can be zinc sulfidesilicon nitride, tantalum oxide, titanium dioxide, or indium tin oxide,for example.

FIG. 53 shows a biosensor cross-sectional profile, in which theone-dimensional grating cross-section is rectangular. Other crosssection profiles of the one dimensional linear grating structure willalso produce the guided mode resonance effect. These include, forexample, triangular or v-shaped, u-shaped, upside-down v- or u-shapes,sinusoidal, trapezoidal, stepped and square. Any regularly repeatingperiodic function will provide a guided mode resonant effect.

Additionally, a one-dimensional linear grating master structure is easyto produce using commercially available gratings, and large-scalegrating master structures with uniform performance can be produced bydeep-ultraviolet (DUV) photolithography. Using sub-micronmicroreplication of a master sensor surface structure on continuoussheets of plastic film, a biosensor can be produced inexpensively overlarge surface areas. A one-dimensional grating biosensor of theinvention can be fabricated by creating a “master” wafer in silicon thatis used as a template for producing the sensor structure on plastic by ahigh-definition microreplication process. The ability to produce ahigh-sensitivity biosensor in plastic over large surface areas enablesincorporation of the biosensor into large area disposable assay formatssuch as microtiter plates and microarray slides. The incorporation of aplastic biosensor into the bottoms, for example, of bottomless 96-wellmicrotiter plates, allows for the use of a biosensor plate to perform,for example, multiple protein-protein binding assays in parallel. Thedetection sensitivity of a plastic-substrate biosensor is equivalent toglass-substrate biosensors. A biosensor structure can incorporated intostandard microtiter plates and used to perform affinity assays based onmeasuring the biochemical interaction between a specific bindingsubstance immobilized on the biosensor surface and binding partnerswithin a test sample. A biosensor can also be incorporated into otherdisposable laboratory assay formats, such as microarray slides, flowcells, and cell culture plates. Incorporation of a biosensor into commonlaboratory formats is desirable for compatibility with existingmicroarray handling equipment such as spotters and incubation chambers.

A one-dimensional linear grating biosensor surface contains an opticalstructure that, when illuminated with collimated white light, isdesigned to reflect only a narrow band of wavelengths. The narrowwavelength band is described as a wavelength “peak.” The “peakwavelength value” (PWV) changes when biological or other material isdeposited or removed from the biosensor surface. A readout instrumentilluminates distinct locations on the biosensor surface with collimatedwhite light, and collects collimated reflected light. The collectedlight is gathered into a wavelength spectrometer for determination ofPWV.

One dimensional linear gratings have resonant characteristics where theilluminating light polarization is oriented perpendicular or parallel tothe grating period. However, a hexagonal grid of holes has betterpolarization symmetry than a rectangular grid of holes. Therefore, acalorimetric resonant reflection biosensor of the invention cancomprise, for example, a two-dimensional hexagonal array of holes (seeFIG. 3B), a two-dimensional array of squares (FIG. 3A) or aone-dimensional grid of parallel lines (see FIG. 2). A one-dimensionallinear grating has the same pitch (i.e. distance between regions of highand low refractive index), period, layer thicknesses, and materialproperties as the hexagonal array grating. However, light must bepolarized perpendicular or parallel to the grating lines in order to beresonantly coupled into the optical structure. Therefore, a polarizingfilter oriented with its polarization axis perpendicular or parallel tothe one-dimensional linear grating must be inserted between theillumination source and the biosensor surface. Because only a smallportion of the illuminating light source is correctly polarized, alonger integration time is required to collect an equivalent amount ofresonantly reflected light compared to a hexagonal grating.

While a one-dimensional linear grating can require either a higherintensity illumination source or a longer measurement integration timecompared to a hexagonal grating, the fabrication requirements for theone-dimensional linear grating structure are simpler. A two-dimensionalhexagonal grating pattern is produced by holographic exposure ofphotoresist to three mutually interfering laser beams. The three beamsare precisely aligned in order to produce a grating pattern that issymmetrical in three directions. A one-dimensional linear gratingpattern requires alignment of only two laser beams to produce aholographic exposure in photoresist, and thus has a reduced alignmentrequirement. A one-dimensional linear grating pattern can also beproduced by, for example, direct writing of photoresist with an electronbeam. Also, several commercially available sources exist for producingone-dimensional linear grating “master” templates for embossing orreplicating a grating structure into plastic. A schematic diagram of alinear grating structure is shown in FIG. 54.

A rectangular grid pattern can be produced in photoresist using anelectron beam direct-write exposure system. A single wafer can beilluminated as a linear grating with two sequential exposures with thepart rotated 90-degrees between exposures.

A one-dimensional or two-dimensional grating can also comprise, forexample, a “stepped” profile, in which high refractive index regions ofa single, fixed height are embedded within a lower refractive indexcover layer. The alternating regions of high and low refractive indexprovide an optical waveguide parallel to the top surface of thebiosensor. See FIG. 5.

For manufacture, a stepped structure is etched or embossed into asubstrate material such as glass or plastic. See FIG. 53B. A uniformthin film of higher refractive index material, such as silicon nitrideor zinc sulfide is deposited on this structure. The deposited layer willfollow the shape contour of the embossed or etched structure in thesubstrate, so that the deposited material has a surface relief profilethat is identical to the original embossed or etched profile. Thethickness of the dielectric layer may be less than, equal to, or greaterthan the depth of the grating structure. The structure can be completedby the application of an optional cover layer comprised of a materialhaving a lower refractive index than the higher refractive indexmaterial and having a substantially flat upper surface. The coveringmaterial can be, for example, glass, epoxy, or plastic.

This structure allows for low cost biosensor manufacturing, because itcan be mass-produced. A “master” grating can be produced in glass,plastic, or metal using, for example, a three-beam laser holographicpatterning process, See e.g., Cowan, The recording and large scaleproduction of crossed holographic grating arrays using multiple beaminterferometry, Proc. Soc. Photo-optical Instum. Eng. 503:120 (1984). Amaster grating can be repeatedly used to emboss a plastic substrate. Theembossed substrate is subsequently coated with a high refractive indexmaterial and optionally, a cover layer.

While a stepped structure is simple to manufacture, it is also possibleto make a resonant biosensor in which the high refractive index materialis not stepped, but which varies with lateral position. FIG. 4 shows aprofile in which the high refractive index material of theone-dimensional or two-dimensional grating, n2, is sinusoidally varyingin height. To produce a resonant reflection at a particular wavelength,the period of the sinusoid is identical to the period of an equivalentstepped structure. The resonant operation of the sinusoidally varyingstructure and its functionality as a biosensor has been verified usingGSOLVER (Grating Solver Development Company, Allen, Tex., USA) computermodels.

Techniques for making two-dimensional gratings are disclosed in Wang, J.Opt. Soc. Am No. 8, August 1990, pp. 1529-44. Biosensors of theinvention can be made in, for example, a semiconductor microfabricationfacility. Biosensors can also be made on a plastic substrate usingcontinuous embossing and optical coating processes. For this type ofmanufacturing process, a “master” structure is built in a rigid materialsuch as glass or silicon, and is used to generate “mother” structures inan epoxy or plastic using one of several types of replicationprocedures. The “mother” structure, in turn, is coated with a thin filmof conducive material, and used as a mold to electroplate a thick filmof nickel. The nickel “daughter” is released from the plastic “mother”structure. Finally, the nickel “daughter” is bonded to a cylindricaldrum, which is used to continuously emboss the surface relief structureinto a plastic film. A device structure that uses an embossed plasticsubstrate is shown in FIG. 5. Following embossing, the plastic structureis overcoated with a thin film of high refractive index material, andoptionally coated with a planarizing, cover layer polymer, and cut toappropriate size.

A substrate for a SWS biosensor can comprise, for example, glass,plastic or epoxy. Optionally, a substrate and a two-dimensional gratingor one-dimensional grating can comprise a single unit. That is, agrating and substrate are formed from the same material, for example,glass, plastic, or epoxy. The surface of a single unit comprising thegrating is coated with a material having a high refractive index, forexample, zinc sulfide, titanium dioxide, tantalum oxide, and siliconnitride. One or more specific binding substances can be immobilized onthe surface of the material having a high refractive index or on anoptional cover layer.

A biosensor of the invention can further comprise a cover layer on thesurface of a two-dimensional grating or one-dimensional grating oppositeof a substrate layer. Where a cover layer is present, the one or morespecific binding substances are immobilized on the surface of the coverlayer opposite of the grating. Preferably, a cover layer comprises amaterial that has a lower refractive index than a material thatcomprises the grating. A cover layer can be comprised of, for example,glass (including spin-on glass (SOG)), epoxy, or plastic.

For example, various polymers that meet the refractive index requirementof a biosensor can be used for a cover layer. SOG can be used due to itsfavorable refractive index, ease of handling, and readiness of beingactivated with specific binding substances using the wealth of glasssurface activation techniques. When the flatness of the biosensorsurface is not an issue for a particular system setup, a gratingstructure of SiN/glass can directly be used as the sensing surface, theactivation of which can be done using the same means as on a glasssurface.

Resonant reflection can also be obtained without a planarizing coverlayer over a two-dimensional grating or one-dimensional grating. Forexample, a biosensor can contain only a substrate coated with astructured thin film layer of high refractive index material. Withoutthe use of a planarizing cover layer, the surrounding medium (such asair or water) fills the grating. Therefore, specific binding substancesare immobilized to the biosensor on all surfaces of a grating exposed tothe specific binding substances, rather than only on an upper surface.

In general, a biosensor of the invention will be illuminated with whitelight that will contain light of every polarization angle. Theorientation of the polarization angle with respect to repeating featuresin a biosensor grating will determine the resonance wavelength. Forexample, a one-dimensional linear grating biosensor structure consistingof a set of repeating lines and spaces will have two opticalpolarizations that can generate separate resonant reflections. Lightthat is polarized perpendicularly to the lines is called “s-polarized,”while light that is polarized parallel to the lines is called“p-polarized.” Both the s and p components of incident light existsimultaneously in an unfiltered illumination beam, and each generates aseparate resonant signal. A biosensor structure can generally bedesigned to optimize the properties of only one polarization (generallythe s-polarization), and the non-optimized polarization is easilyremoved by a polarizing filter.

In order to remove the polarization dependence, so that everypolarization angle generates the same resonant reflection spectra, analternate biosensor structure can be used that consists of a set ofconcentric rings. In this structure, the difference between the insidediameter and the outside diameter of each concentric ring is equal toabout one-half of a grating period. Each successive ring has an insidediameter that is about one grating period greater than the insidediameter of the previous ring. The concentric ring pattern extends tocover a single sensor location—such as a microarray spot or a microtiterplate well. Each separate microarray spot or microtiter plate well has aseparate concentric ring pattern centered within it. e.g., FIG. 49. Allpolarization directions of such a structure have the samecross-sectional profile. The concentric ring structure must beilluminated precisely on-center to preserve polarization independence.The grating period of a concentric ring structure is less than thewavelength of the resonantly reflected light. The grating period isabout 0.01 micron to about 1 micron. The grating depth is about 0.01 toabout 1 micron.

In another embodiment, an array of holes or posts are arranged toclosely approximate the concentric circle structure described abovewithout requiring the illumination beam to be centered upon anyparticular location of the grid. See e.g. FIG. 50. Such an array patternis automatically generated by the optical interference of three laserbeams incident on a surface from three directions at equal angles. Inthis pattern, the holes (or posts) are centered upon the corners of anarray of closely packed hexagons as shown in FIG. 50. The holes or postsalso occur in the center of each hexagon. Such a hexagonal grid of holesor posts has three polarization directions that “see” the samecross-sectional profile. The hexagonal grid structure, therefore,provides equivalent resonant reflection spectra using light of anypolarization angle. Thus, no polarizing filter is required to removeunwanted reflected signal components. The period of the holes or postscan be about 0.01 microns to about 1 micron and the depth or height canbe about 0.01 microns to about 1 micron.

The invention provides a resonant reflection structures and transmissionfilter structures comprising concentric circle gratings and hexagonalgrids of holes or posts. For a resonant reflection structure, lightoutput is measured on the same side of the structure as the illuminatinglight beam. For a transmission filter structure, light output ismeasured on the opposite side of the structure as the illuminating beam.The reflected and transmitted signals are complementary. That is, if awavelength is strongly reflected, it is weakly transmitted. Assuming noenergy is absorbed in the structure itself, the reflected+transmittedenergy at any given wavelength is constant. The resonant reflectionstructure and transmission filters are designed to give a highlyefficient reflection at a specified wavelength. Thus, a reflectionfilter will “pass” a narrow band of wavelengths, while a transmissionfilter will “cut” a narrow band of wavelengths from incident light.

A resonant reflection structure or a transmission filter structure cancomprise a two-dimensional grating arranged in a pattern of concentriccircles. A resonant reflection structure or transmission filterstructure can also comprise a hexagonal grid of holes or posts. Whenthese structure are illuminated with an illuminating light beam, areflected radiation spectrum is produced that is independent of anillumination polarization angle of the illuminating light beam. Whenthese structures are illuminated a resonant grating effect is producedon the reflected radiation spectrum, wherein the depth and period of thetwo-dimensional grating or hexagonal grid of holes or posts are lessthan the wavelength of the resonant grating effect. These structuresreflect a narrow band of light when the structure is illuminated with abroadband of light.

Resonant reflection structures and transmission filter structures of theinvention can be used as biosensors. For example, one or more specificbinding substances can be immobilized on the hexagonal grid of holes orposts or on the two-dimensional grating arranged in concentric circles.

In one embodiment of the invention, a reference resonant signal isprovided for more accurate measurement of peak resonant wavelengthshifts. The reference resonant signal can cancel out environmentaleffects, including, for example, temperature. A reference signal can beprovided using a resonant reflection superstructure that produces twoseparate resonant wavelengths. A transparent resonant reflectionsuperstructure can contain two sub-structures. A first sub-structurecomprises a first one- or two-dimensional grating with a top and abottom surface. The top surface of a one- or two-dimensional gratingcomprises the grating surface. The first one- or two-dimensional gratingcan comprise one or more specific binding substances immobilized on itstop surface. The top surface of the first one- or two-dimensionalgrating is in contact with a test sample. An optional substrate layercan be present to support the bottom surface of the first one- ortwo-dimensional grating. The substrate layer comprises a top and bottomsurface. The top surface of the substrate is in contact with, andsupports the bottom surface of the first one- or two-dimensionalgrating.

A second sub-structure comprises a second one- or two-dimensionalgrating with a top surface and a bottom surface. The second one- ortwo-dimensional grating is not in contact with a test sample. The secondone- or two-dimensional grating can be fabricated onto the bottomsurface of the substrate that supports the first one- or two-dimensionalgrating. Where the second one- or two-dimensional grating is fabricatedon the substrate that supports the first one- or two-dimensionalgrating, the bottom surface of the second one- or two-dimensionalgrating can be fabricated onto the bottom surface of the substrate.Therefore, the top surface of the second one- or two-dimensional gratingwill face the opposite direction of the top surface of the first one- ortwo-dimensional grating.

The top surface of the second one- or two-dimensional grating can alsobe attached directly to the bottom surface of the first sub-structure.In this embodiment the top surface of the second one- or two-dimensionalgrating will face the same direction as the top surface of the firstone- or two-dimensional grating. A substrate can support the bottomsurface of the second one- or two-dimensional grating in thisembodiment.

Because the second sub-structure is not in physical contact with thetest sample, its peak resonant wavelength is not subject to changes inthe optical density of the test media, or deposition of specific bindingsubstances or binding partners on the surface of the first one- ortwo-dimensional grating. Therefore, such a superstructure produces tworesonant signals. Because the location of the peak resonant wavelengthin the second sub-structure is fixed, the difference in peak resonantwavelength between the two sub-structures provides a relative means fordetermining the amount of specific binding substances or bindingpartners or both deposited on the top surface of the first substructurethat is exposed to the test sample.

A biosensor superstructure can be illuminated from its top surface orfrom its bottom surface, or from both surfaces. The peak resonancereflection wavelength of the first substructure is dependent on theoptical density of material in contact with the superstructure surface,while the peak resonance reflection wavelength of the secondsubstructure is independent of the optical density of material incontact with the superstructure surface.

In one embodiment of the invention, a biosensor is illuminated from thebottom surface of the biosensor. Approximately 50% of the incident lightis reflected from the bottom surface of biosensor without reaching theactive (top) surface of the biosensor. A thin film or physical structurecan be included in a biosensor composition that is capable of maximizingthe amount of light that is transmitted to the upper surface of thebiosensor while minimizing the reflected energy at the resonantwavelength. The anti-reflection thin film or physical structure of thebottom surface of the biosensor can comprise, for example, a singledielectric thin film, a stack of multiple dielectric thin films, or a“motheye” structure that is embossed into the bottom biosensor surface.An example of a motheye structure is disclosed in Hobbs, et al.“Automated interference lithography system for generation of sub-micronfeature size patterns,” Proc. 1999 Micromachine Technology forDiffracting and Holographic Optics, Society of Photo-OpticalInstrumentation Engineers, p. 124-135, (1999).

In one embodiment of the invention, an optical device is provided. Anoptical device comprises a structure similar to any biosensor of theinvention; however, an optical device does not comprise one of morebinding substances immobilized on the two-dimensional grating. Anoptical device can be used as a narrow band optical filter.

In one embodiment of the invention, an interaction of a first moleculewith a second test molecule can be detected. A SWS biosensor asdescribed above is used; however, there are no specific bindingsubstances immobilized on its surface. Therefore, the biosensorcomprises a one- or two-dimensional grating, a substrate layer thatsupports the one- or two-dimensional grating, and optionally, a coverlayer. As described above, when the biosensor is illuminated a resonantgrating effect is produced on the reflected radiation spectrum, and thedepth and period of the grating are less than the wavelength of theresonant grating effect.

To detect an interaction of a first molecule with a second testmolecule, a mixture of the first and second molecules is applied to adistinct location on a biosensor. A distinct location can be one spot orwell on a biosensor or can be a large area on a biosensor. A mixture ofthe first molecule with a third control molecule is also applied to adistinct location on a biosensor. The biosensor can be the samebiosensor as described above, or can be a second biosensor. If thebiosensor is the same biosensor, a second distinct location can be usedfor the mixture of the first molecule and the third control molecule.Alternatively, the same distinct biosensor location can be used afterthe first and second molecules are washed from the biosensor. The thirdcontrol molecule does not interact with the first molecule and is aboutthe same size as the first molecule. A shift in the reflected wavelengthof light from the distinct locations of the biosensor or biosensors ismeasured. If the shift in the reflected wavelength of light from thedistinct location having the first molecule and the second test moleculeis greater than the shift in the reflected wavelength from the distinctlocation having the first molecule and the third control molecule, thenthe first molecule and the second test molecule interact. Interactioncan be, for example, hybridization of nucleic acid molecules, specificbinding of an antibody or antibody fragment to an antigen, and bindingof polypeptides. A first molecule, second test molecule, or thirdcontrol molecule can be, for example, a nucleic acid, polypeptide,antigen, polyclonal antibody, monoclonal antibody, single chain antibody(scFv), F(ab) fragment, F(ab′)₂ fragment, Fv fragment, small organicmolecule, cell, virus, and bacteria.

Specific Binding Substances and Binding Partners

One or more specific binding substances are immobilized on the one- ortwo-dimensional grating or cover layer, if present, by for example,physical adsorption or by chemical binding. A specific binding substancecan be, for example, a nucleic acid, polypeptide, antigen, polyclonalantibody, monoclonal antibody, single chain antibody (scFv), F(ab)fragment, F(ab′)₂ fragment, Fv fragment, small organic molecule, cell,virus, bacteria, polymer, peptide solutions, single- or double-strandedDNA solutions, RNA solutions, solutions containing compounds from acombinatorial chemical library, or biological sample. A biologicalsample can be for example, blood, plasma, serum, gastrointestinalsecretions, homogenates of tissues or tumors, synovial fluid, feces,saliva, sputum, cyst fluid, amniotic fluid, cerebrospinal fluid,peritoneal fluid, lung lavage fluid, semen, lymphatic fluid, tears, orprostatitc fluid.

Preferably, one or more specific binding substances are arranged in amicroarray of distinct locations on a biosensor. A microarray ofspecific binding substances comprises one or more specific bindingsubstances on a surface of a biosensor of the invention such that asurface contains many distinct locations, each with a different specificbinding substance or with a different amount of a specific bindingsubstance. For example, an array can comprise 1, 10, 100, 1,000, 10,000,or 100,000 distinct locations. Such a biosensor surface is called amicroarray because one or more specific binding substances are typicallylaid out in a regular grid pattern in x-y coordinates. However, amicroarray of the invention can comprise one or more specific bindingsubstance laid out in any type of regular or irregular pattern. Forexample, distinct locations can define a microarray of spots of one ormore specific binding substances. A microarray spot can be about 50 toabout 500 microns in diameter. A microarray spot can also be about 150to about 200 microns in diameter. One or more specific bindingsubstances can be bound to their specific binding partners.

A microarray on a biosensor of the invention can be created by placingmicrodroplets of one or more specific binding substances onto, forexample, an x-y grid of locations on a one- or two-dimensional gratingor cover layer surface. When the biosensor is exposed to a test samplecomprising one or more binding partners, the binding partners will bepreferentially attracted to distinct locations on the microarray thatcomprise specific binding substances that have high affinity for thebinding partners. Some of the distinct locations will gather bindingpartners onto their surface, while other locations will not.

A specific binding substance specifically binds to a binding partnerthat is added to the surface of a biosensor of the invention. A specificbinding substance specifically binds to its binding partner, but doesnot substantially bind other binding partners added to the surface of abiosensor. For example, where the specific binding substance is anantibody and its binding partner is a particular antigen, the antibodyspecifically binds to the particular antigen, but does not substantiallybind other antigens. A binding partner can be, for example, a nucleicacid, polypeptide, antigen, polyclonal antibody, monoclonal antibody,single chain antibody (scFv), F(ab) fragment, F(ab′)₂ fragment, Fvfragment, small organic molecule, cell, virus, bacteria, polymer,peptide solutions, single- or double-stranded DNA solutions, RNAsolutions, solutions containing compounds from a combinatorial chemicallibrary and biological sample. A biological sample can be, for example,blood, plasma, serum, gastrointestinal secretions, homogenates oftissues or tumors, synovial fluid, feces, saliva, sputum, cyst fluid,amniotic fluid, cerebrospinal fluid, peritoneal fluid, lung lavagefluid, semen, lymphatic fluid, tears, and prostatitc fluid.

One example of a microarray of the invention is a nucleic acidmicroarray, in which each distinct location within the array contains adifferent nucleic acid molecule. In this embodiment, the spots withinthe nucleic acid microarray detect complementary chemical binding withan opposing strand of a nucleic acid in a test sample.

While microtiter plates are the most common format used for biochemicalassays, microarrays are increasingly seen as a means for maximizing thenumber of biochemical interactions that can be measured at one timewhile minimizing the volume of precious reagents. By application ofspecific binding substances with a microarray spotter onto a biosensorof the invention, specific binding substance densities of 10,000specific binding substances/in² can be obtained. By focusing anillumination beam to interrogate a single microarray location, abiosensor can be used as a label-free microarray readout system.

Immobilization or One or More Specific Binding Substances

Immobilization of one or more binding substances onto a biosensor isperformed so that a specific binding substance will not be washed awayby rinsing procedures, and so that its binding to binding partners in atest sample is unimpeded by the biosensor surface. Several differenttypes of surface chemistry strategies have been implemented for covalentattachment of specific binding substances to, for example, glass for usein various types of microarrays and biosensors. These same methods canbe readily adapted to a biosensor of the invention. Surface preparationof a biosensor so that it contains the correct functional groups forbinding one or more specific binding substances is an integral part ofthe biosensor manufacturing process.

One or more specific binding substances can be attached to a biosensorsurface by physical adsorption (i.e., without the use of chemicallinkers) or by chemical binding (i.e., with the use of chemicallinkers). Chemical binding can generate stronger attachment of specificbinding substances on a biosensor surface and provide definedorientation and conformation of the surface-bound molecules.

Several examples of chemical binding of specific binding substances to abiosensor of the invention appear in Example 8, below. Other types ofchemical binding include, for example, amine activation, aldehydeactivation, and nickel activation. These surfaces can be used to attachseveral different types of chemical linkers to a biosensor surface, asshown in FIG. 6. While an amine surface can be used to attach severaltypes of linker molecules, an aldehyde surface can be used to bindproteins directly, without an additional linker. A nickel surface can beused to bind molecules that have an incorporated histidine (“his”) tag.Detection of “his-tagged” molecules with a nickel-activated surface iswell known in the art (Whitesides, Anal Chem. 68, 490, (1996)).

Immobilization of specific binding substances to plastic, epoxy, or highrefractive index material can be performed essentially as described forimmobilization to glass. However, the acid wash step can be eliminatedwhere such a treatment would damage the material to which the specificbinding substances are immobilized.

For the detection of binding partners at concentrations less than about˜0.1 ng/ml, it is preferable to amplify and transduce binding partnersbound to a biosensor into an additional layer on the biosensor surface.The increased mass deposited on the biosensor can be easily detected asa consequence of increased optical path length. By incorporating greatermass onto a biosensor surface, the optical density of binding partnerson the surface is also increased, thus rendering a greater resonantwavelength shift than would occur without the added mass. The additionof mass can be accomplished, for example, enzymatically, through a“sandwich” assay, or by direct application of mass to the biosensorsurface in the form of appropriately conjugated beads or polymers ofvarious size and composition. This principle has been exploited forother types of optical biosensors to demonstrate sensitivity increasesover 1500× beyond sensitivity limits achieved without massamplification. See, e.g., Jenison et al., “Interference-based detectionof nucleic acid targets on optically coated silicon,” NatureBiotechnology, 19: 62-65, 2001.

As an example, FIG. 7A shows that an NH₂-activated biosensor surface canhave a specific binding substance comprising a single-strand DNA captureprobe immobilized on the surface. The capture probe interactsselectively with its complementary target binding partner. The bindingpartner, in turn, can be designed to include a sequence or tag that willbind a “detector” molecule. As shown in FIG. 7A, a detector molecule cancontain, for example, a linker to horseradish peroxidase (HRP) that,when exposed to the correct enzyme, will selectively deposit additionalmaterial on the biosensor only where the detector molecule is present.Such a procedure can add, for example, 300 angstroms of detectablebiomaterial to the biosensor within a few minutes.

A “sandwich” approach can also be used to enhance detection sensitivity.In this approach, a large molecular weight molecule can be used toamplify the presence of a low molecular weight molecule. For example, abinding partner with a molecular weight of, for example, about 0.1 kDato about 20 kDa, can be tagged with, for example,succinimidyl-6-[a-methyl-a-(2-pyridyl-dithio)toluamido]hexanoate (SMPT),or dimethylpimelimidate (DMP), histidine, or a biotin molecule, as shownin FIG. 7B. Where the tag is biotin, the biotin molecule will bindsstrongly with streptavidin, which has a molecular weight of 60 kDa.Because the biotin/streptavidin interaction is highly specific, thestreptavidin amplifies the signal that would be produced only by thesmall binding partner by a factor of 60.

Detection sensitivity can be further enhanced through the use ofchemically derivatized small particles. “Nanoparticles” made ofcolloidal gold, various plastics, or glass with diameters of about 3-300nm can be coated with molecular species that will enable them tocovalently bind selectively to a binding partner. For example, as shownin FIG. 7C, nanoparticles that are covalently coated with streptavidincan be used to enhance the visibility of biotin-tagged binding partnerson the biosensor surface. While a streptavidin molecule itself has amolecular weight of 60 kDa, the derivatized bead can have a molecularweight of any size, including, for example, 60 KDa. Binding of a largebead will result in a large change in the optical density upon thebiosensor surface, and an easily measurable signal. This method canresult in an approximately 1000× enhancement in sensitivity resolution.

Surface-Relief Volume Diffractive Biosensors

Another embodiment of the invention is a biosensor that comprises volumesurface-relief volume diffractive structures (a SRVD biosensor). SRVDbiosensors have a surface that reflect predominantly at a particularnarrow band of optical wavelengths when illuminated with a broad band ofoptical wavelengths. Where specific binding substances and/or bindingpartners are immobilized on a SRVD biosensor, the reflected wavelengthof light is shifted. One-dimensional surfaces, such as thin filminterference filters and Bragg reflectors, can select a narrow range ofreflected or transmitted wavelengths from a broadband excitation source,however, the deposition of additional material, such as specific bindingsubstances and/or binding partners onto their upper surface results onlyin a change in the resonance linewidth, rather than the resonancewavelength. In contrast, SRVD biosensors have the ability to alter thereflected wavelength with the addition of material, such as specificbinding substances and/or binding partners to the surface.

A SRVD biosensor comprises a sheet material having a first and secondsurface. The first surface of the sheet material defines relief volumediffraction structures. A sheet material can be comprised of, forexample, plastic, glass, semiconductor wafer, or metal film.

A relief volume diffractive structure can be, for example, atwo-dimensional grating, as described above, or a three-dimensionalsurface-relief volume diffractive grating. The depth and period ofrelief volume diffraction structures are less than the resonancewavelength of light reflected from a biosensor.

A three-dimensional surface-relief volume diffractive grating can be,for example, a three-dimensional phase-quantized terraced surface reliefpattern whose groove pattern resembles a stepped pyramid. When such agrating is illuminated by a beam of broadband radiation, light will becoherently reflected from the equally spaced terraces at a wavelengthgiven by twice the step spacing times the index of refraction of thesurrounding medium. Light of a given wavelength is resonantly diffractedor reflected from the steps that are a half-wavelength apart, and with abandwidth that is inversely proportional to the number of steps. Thereflected or diffracted color can be controlled by the deposition of adielectric layer so that a new wavelength is selected, depending on theindex of refraction of the coating.

A stepped-phase structure can be produced first in photoresist bycoherently exposing a thin photoresist film to three laser beams, asdescribed previously. See e.g., Cowen, “The recording and large scalereplication of crossed holographic grating arrays using multiple beaminterferometry,” in International Conference on the Application, Theory,and Fabrication of Periodic Structures, Diffraction Gratings, and MoirePhenomena II, Lerner, ed., Proc. Soc. Photo-Opt. Instrum. Eng., 503,120-129, 1984; Cowen, “Holographic honeycomb microlens,” Opt. Eng. 24,796-802 (1985); Cowen & Slafer, “The recording and replication ofholographic micropatterns for the ordering of photographic emulsiongrains in film systems,” J. Imaging Sci. 31, 100-107, 1987. Thenonlinear etching characteristics of photoresist are used to develop theexposed film to create a three-dimensional relief pattern. Thephotoresist structure is then replicated using standard embossingprocedures. For example, a thin silver film is deposited over thephotoresist structure to form a conducting layer upon which a thick filmof nickel can be electroplated. The nickel “master” plate is then usedto emboss directly into a plastic film, such as vinyl, that has beensoftened by heating or solvent.

The theory describing the design and fabrication of three-dimensionalphase-quantized terraced surface relief pattern that resemble steppedpyramids is described: Cowen, “Aztec surface-relief volume diffractivestructure,” J. Opt. Soc. Am. A, 7:1529 (1990).

An example of a three-dimensional phase-quantized terraced surfacerelief pattern is a pattern that resembles a stepped pyramid. Eachinverted pyramid is approximately 1 micron in diameter, preferably, eachinverted pyramid can be about 0.5 to about 5 microns diameter, includingfor example, about 1 micron. The pyramid structures can be close-packedso that a typical microarray spot with a diameter of 150-200 microns canincorporate several hundred stepped pyramid structures. The reliefvolume diffraction structures have a period of about 0.1 to about 1micron and a depth of about 0.1 to about 1 micron. FIG. 8 demonstrateshow individual microarray locations (with an entire microarray spotincorporating hundreds of pyramids now represented by a single pyramidfor one microarray spot) can be optically queried to determine ifspecific binding substances or binding partners are adsorbed onto thesurface. When the structure is illuminated with white light, structureswithout significant bound material will reflect wavelengths determinedby the step height of the structure. When higher refractive indexmaterial, such as binding partners or specific binding substances, areincorporated over the reflective metal surface, the reflected wavelengthis modified to shift toward longer wavelengths. The color that isreflected from the terraced step structure is theoretically given astwice the step height times the index of refraction of a reflectivematerial that is coated onto the first surface of a sheet material of aSRVD biosensor. A reflective material can be, for example silver,aluminum, or gold.

One or more specific binding substances, as described above, areimmobilized on the reflective material of a SRVD biosensor. One or morespecific binding substances can be arranged in microarray of distinctlocations, as described above, on the reflective material. FIG. 9provides an example of a 9-element microarray biosensor. Many individualgrating structures, represented by small circles, lie within eachmicroarray spot. The microarray spots, represented by the largercircles, will reflect white light in air at a wavelength that isdetermined by the refractive index of material on their surface.Microarray locations with additional adsorbed material will havereflected wavelengths that are shifted toward longer wavelengths,represented by the larger circles.

Because the reflected wavelength of light from a SRVD biosensor isconfined to a narrow bandwidth, very small changes in the opticalcharacteristics of the surface manifest themselves in easily observedchanges in reflected wavelength spectra. The narrow reflection bandwidthprovides a surface adsorption sensitivity advantage compared toreflectance spectrometry on a flat surface.

A SRVD biosensor reflects light predominantly at a first single opticalwavelength when illuminated with a broad band of optical wavelengths,and reflects light at a second single optical wavelength when one ormore specific binding substances are immobilized on the reflectivesurface. The reflection at the second optical wavelength results fromoptical interference. A SRVD biosensor also reflects light at a thirdsingle optical wavelength when the one or more specific bindingsubstances are bound to their respective binding partners, due tooptical interference.

Readout of the reflected color can be performed serially by focusing amicroscope objective onto individual microarray spots and reading thereflected spectrum, or in parallel by, for example, projecting thereflected image of the microarray onto a high resolution color CCDcamera.

A SRVD biosensor can be manufactured by, for example, producing a metalmaster plate, and stamping a relief volume diffractive structure into,for example, a plastic material like vinyl. After stamping, the surfaceis made reflective by blanket deposition of, for example, a thin metalfilm such as gold, silver, or aluminum. Compared to MEMS-basedbiosensors that rely upon photolithography, etching, and wafer bondingprocedures, the manufacture of a SRVD biosensor is very inexpensive.

Liquid-Containing Vessels

A SWS or SRVD biosensor of the invention can comprise an inner surface,for example, a bottom surface of a liquid-containing vessel. Aliquid-containing vessel can be, for example, a microtiter plate well, atest tube, a petri dish, or a microfluidic channel. One embodiment ofthis invention is a SWS or SRVD biosensor that is incorporated into anytype of microtiter plate. For example, a SWS biosensor or SRVD biosensorcan be incorporated into the bottom surface of a microtiter plate byassembling the walls of the reaction vessels over the resonantreflection surface, as shown in FIG. 10, so that each reaction “spot”can be exposed to a distinct test sample. Therefore, each individualmicrotiter plate well can act as a separate reaction vessel. Separatechemical reactions can, therefore, occur within adjacent wells withoutintermixing reaction fluids and chemically distinct test solutions canbe applied to individual wells.

Several methods for attaching a biosensor of the invention to the bottomsurface of bottomless microtiter plates can be used, including, forexample, adhesive attachment, ultrasonic welding, and laser welding.

The most common assay formats for pharmaceutical high-throughputscreening laboratories, molecular biology research laboratories, anddiagnostic assay laboratories are microtiter plates. The plates arestandard-sized plastic cartridges that can contain 96, 384, or 1536individual reaction vessels arranged in a grid. Due to the standardmechanical configuration of these plates, liquid dispensing, roboticplate handling, and detection systems are designed to work with thiscommon format. A biosensor of the invention can be incorporated into thebottom surface of a standard microtiter plate. See, e.g., FIG. 10.Because the biosensor surface can be fabricated in large areas, andbecause the readout system does not make physical contact with thebiosensor surface, an arbitrary number of individual biosensor areas canbe defined that are only limited by the focus resolution of theillumination optics and the x-y stage that scans theillumination/detection probe across the biosensor surface.

Holding Fixtures

Any number of biosensors that are, for example, about 1 mm² to about 5mm², and preferably less than about 3×3 mm² can be arranged onto aholding fixture that can simultaneously dip the biosensors into separateliquid-containing vessels, such as wells of a microtiter plate, forexample, a 96-, 384-, or 1536-well microtiter plate. See e.g., FIG. 11.Each of the biosensors can contain multiple distinct locations. Aholding fixture has one or more biosensors attached to the holdingfixture so that each individual biosensor can be lowered into a separateliquid-containing vessel. A holding fixture can comprise plastic, epoxyor metal. For example, 50, 96, 384, or 1,000, or 1,536 biosensors can bearranged on a holding fixture, where each biosensor has 25, 100, 500, or1,000 distinct locations. As an example, where 96 biosenors are attachedto a holding fixture and each biosensor comprises 100 distinctlocations, 9600 biochemical assays can be performed simultaneously.

Methods of Using SWS and SRVD Biosensors

SWS and SRVD biosensors of the invention can be used to study one or anumber of specific binding substance/binding partner interactions inparallel. Binding of one or more specific binding substances to theirrespective binding partners can be detected, without the use of labels,by applying one or more binding partners to a SWS or SRVD biosensor thathave one or more specific binding substances immobilized on theirsurfaces. A SWS biosensor is illuminated with light and a maxima inreflected wavelength, or a minima in transmitted wavelength of light isdetected from the biosensor. If one or more specific binding substanceshave bound to their respective binding partners, then the reflectedwavelength of light is shifted as compared to a situation where one ormore specific binding substances have not bound to their respectivebinding partners. Where a SWS biosensor is coated with an array ofdistinct locations containing the one or more specific bindingsubstances, then a maxima in reflected wavelength or minima intransmitted wavelength of light is detected from each distinct locationof the biosensor.

A SRVD biosensor is illuminated with light after binding partners havebeen added and the reflected wavelength of light is detected from thebiosensor. Where one or more specific binding substances have bound totheir respective binding partners, the reflected wavelength of light isshifted.

In one embodiment of the invention, a variety of specific bindingsubstances, for example, antibodies, can be immobilized in an arrayformat onto a biosensor of the invention. The biosensor is thencontacted with a test sample of interest comprising binding partners,such as proteins. Only the proteins that specifically bind to theantibodies immobilized on the biosensor remain bound to the biosensor.Such an approach is essentially a large-scale version of anenzyme-linked immunosorbent assay; however, the use of an enzyme orfluorescent label is not required.

The activity of an enzyme can be detected by applying one or moreenzymes to a SWS or SRVD biosensor to which one or more specific bindingsubstances have been immobilized. The biosensor is washed andilluminated with light. The reflected wavelength of light is detectedfrom the biosensor. Where the one or more enzymes have altered the oneor more specific binding substances of the biosensor by enzymaticactivity, the reflected wavelength of light is shifted.

Additionally, a test sample, for example, cell lysates containingbinding partners can be applied to a biosensor of the invention,followed by washing to remove unbound material. The binding partnersthat bind to a biosensor can be eluted from the biosensor and identifiedby, for example, mass spectrometry. Optionally, a phage DNA displaylibrary can be applied to a biosensor of the invention followed bywashing to remove unbound material. Individual phage particles bound tothe biosensor can be isolated and the inserts in these phage particlescan then be sequenced to determine the identity of the binding partner.

For the above applications, and in particular proteomics applications,the ability to selectively bind material, such as binding partners froma test sample onto a biosensor of the invention, followed by the abilityto selectively remove bound material from a distinct location of thebiosensor for further analysis is advantageous. Biosensors of theinvention are also capable of detecting and quantifying the amount of abinding partner from a sample that is bound to a biosensor arraydistinct location by measuring the shift in reflected wavelength oflight. For example, the wavelength shift at one distinct biosensorlocation can be compared to positive and negative controls at otherdistinct biosensor locations to determine the amount of a bindingpartner that is bound to a biosensor array distinct location.

SWS and Electrically Conducting Material

An optional biosensor structure can further enable a biosensor array toselectively attract or repel binding partners from individual distinctlocations on a biosensor. As is well known in the art, an electromotiveforce can be applied to biological molecules such as nucleic acids andamino acids subjecting them to an electric field. Because thesemolecules are electronegative, they are attracted to a positivelycharged electrode and repelled by a negatively charged electrode.

A grating structure of a resonant optical biosensor can be built usingan electrically conducting material rather than an electricallyinsulating material. An electric field can be applied near the biosensorsurface. Where a grating operates as both a resonant reflector biosensorand as an electrode, the grating comprises a material that is bothoptically transparent near the resonant wavelength, and has lowresistivity. In one embodiment of the invention, the material is indiumtin oxide, InSn_(x)O_(1-x) (ITO). ITO is commonly used to producetransparent electrodes for flat panel optical displays, and is thereforereadily available at low cost on large glass sheets. The refractiveindex of ITO can be adjusted by controlling x, the fraction of Sn thatis present in the material. Because the liquid test sample solution willhave mobile ions (and will therefore be an electrical conductor) it isnecessary for the ITO electrodes to be coated with an insulatingmaterial. For the resonant optical biosensor, a grating layer is coatedwith a layer with lower refractive index material. Materials such ascured photoresist (n=1.65), cured optical epoxy (n=1.5), and glass(n=1.4-1.5) are strong electrical insulators that also have a refractiveindex that is lower than ITO (n=2.0-2.65). A cross-sectional diagram ofa biosensor that incorporates an ITO grating is shown in FIG. 48. n₁represents the refractive index of an electrical insulator. n₂represents the refractive index of a two-dimensional grating. t₁represents the thickness of the electrical insulator. t₂ represents thethickness of the two-dimensional grating. n_(bio) represents therefractive index of one or more specific binding substances and t_(BIO)represents the thickness of the one or more specific binding substances.

A grating can be a continuous sheet of ITO that contains an array ofregularly spaced holes. The holes are filled in with an electricallyinsulating material, such as cured photoresist. The electricallyinsulating layer overcoats the ITO grating so that the upper surface ofthe structure is completely covered with electrical insulator, and sothat the upper surface is substantially flat. When the biosensor isilluminated with light a resonant grating effect is produced on thereflected radiation spectrum. The depth and the period of the gratingare less than the wavelength of the resonant grating effect.

As shown in FIG. 12 and FIG. 13, a single electrode can comprise aregion that contains many grating periods. Building two or more separategrating regions on the same substrate surface creates an array ofbiosensor electrodes. Electrical contact to each biosensor electrode isprovided using an electrically conducting trace that is built from thesame material as the conductor within the biosensor electrode. Theconducting trace is connected to a voltage source that can apply anelectrical potential to the electrode. To apply an electrical potentialto the biosensor that is capable of attracting or repelling a moleculenear the electrode surface, a biosensor upper surface can be immersed ina liquid sample as shown in FIG. 14. A “common” electrode can be placedwithin the sample liquid, and a voltage can be applied between oneselected biosensor electrode region and the common electrode. In thisway, one, several, or all electrodes can be activated or inactivated ata given time. FIG. 15 illustrates the attraction of electronegativemolecules to the biosensor surface when a positive voltage is applied tothe electrode, while FIG. 16 illustrates the application of a repellingforce such as a reversed electrical charge to electronegative moleculesusing a negative electrode voltage.

Detection Systems

A detection system can comprise a biosensor of the invention, a lightsource that directs light to the biosensor, and a detector that detectslight reflected from the biosensor. In one embodiment, it is possible tosimplify the readout instrumentation by the application of a filter sothat only positive results over a determined threshold trigger adetection.

A light source can illuminate a biosensor from its top surface, i.e.,the surface to which one or more specific binding substances areimmobilized or from its bottom surface. By measuring the shift inresonant wavelength at each distinct location of a biosensor of theinvention, it is possible to determine which distinct locations havebinding partners bound to them. The extent of the shift can be used todetermine the amount of binding partners in a test sample and thechemical affinity between one or more specific binding substances andthe binding partners of the test sample.

A biosensor of the invention can be illuminated twice. The firstmeasurement determines the reflectance spectra of one or more distinctlocations of a biosensor array with one or more specific bindingsubstances immobilized on the biosensor. The second measurementdetermines the reflectance spectra after one or more binding partnersare applied to a biosensor. The difference in peak wavelength betweenthese two measurements is a measurement of the amount of bindingpartners that have specifically bound to a biosensor or one or moredistinct locations of a biosensor. This method of illumination cancontrol for small nonuniformities in a surface of a biosensor that canresult in regions with slight variations in the peak resonantwavelength. This method can also control for varying concentrations ormolecular weights of specific binding substances immobilized on abiosensor

Computer simulation can be used to determine the expected dependencebetween a peak resonance wavelength and an angle of incidentillumination. A biosensor structure as shown in FIG. 1 can be forpurposes of demonstration. The substrate chosen was glass(n_(substrate)=1.50). The grating is a two-dimensional pattern ofsilicon nitride squares (t₂=180 nm, n₂=2.01 (n=refractive index),k₂=0.001 (k=absorption coefficient)) with a period of 510 nm, and afilling factor of 56.2% (i.e., 56.2% of the surface is covered withsilicon nitride squares while the rest is the area between the squares).The areas between silicon nitride squares are filled with a lowerrefractive index material. The same material also covers the squares andprovides a uniformly flat upper surface. For this simulation, a glasslayer was selected (n₁=1.40) that covers the silicon nitride squares byt₂=100 nm.

The reflected intensity as a function of wavelength was modeled usingGSOLVER software, which utilizes full 3-dimensional vector code usinghybrid Rigorous Coupled Wave Analysis and Modal analysis. GSOLVERcalculates diffracted fields and diffraction efficiencies from planewave illumination of arbitrarily complex grating structures. Theillumination can be from any incidence and any polarization.

FIG. 19 plots the dependence of the peak resonant wavelength upon theincident illumination angle. The simulation shows that there is a strongcorrelation between the angle of incident light, and the peak wavelengththat is measured. This result implies that the collimation of theilluminating beam, and the alignment between the illuminating beam andthe reflected beam will directly affect the resonant peak linewidth thatis measured. If the collimation of the illuminating beam is poor, arange illuminating angles will be incident on the biosensor surface, anda wider resonant peak will be measured than if purely collimated lightwere incident.

Because the lower sensitivity limit of a biosensor is related to theability to determine the peak maxima, it is important to measure anarrow resonant peak. Therefore, the use of a collimating illuminationsystem with the biosensor provides for the highest possible sensitivity.

One type of detection system for illuminating the biosensor surface andfor collecting the reflected light is a probe containing, for example,six illuminating optical fibers that are connected to a light source,and a single collecting optical fiber connected to a spectrometer. Thenumber of fibers is not critical, any number of illuminating orcollecting fibers are possible. The fibers are arranged in a bundle sothat the collecting fiber is in the center of the bundle, and issurrounded by the six illuminating fibers. The tip of the fiber bundleis connected to a collimating lens that focuses the illumination ontothe surface of the biosensor.

In this probe arrangement, the illuminating and collecting fibers areside-by-side. Therefore, when the collimating lens is correctly adjustedto focus light onto the biosensor surface, one observes six clearlydefined circular regions of illumination, and a central dark region.Because the biosensor does not scatter light, but rather reflects acollimated beam, no light is incident upon the collecting fiber, and noresonant signal is observed. Only by defocusing the collimating lensuntil the six illumination regions overlap into the central region isany light reflected into the collecting fiber. Because only defocused,slightly uncollimated light can produce a signal, the biosensor is notilluminated with a single angle of incidence, but with a range ofincident angles. The range of incident angles results in a mixture ofresonant wavelengths due to the dependence shown in FIG. 19. Thus, widerresonant peaks are measured than might otherwise be possible.

Therefore, it is desirable for the illuminating and collecting fiberprobes to spatially share the same optical path. Several methods can beused to co-locate the illuminating and collecting optical paths. Forexample, a single illuminating fiber, which is connected at its firstend to a light source that directs light at the biosensor, and a singlecollecting fiber, which is connected at its first end to a detector thatdetects light reflected from the biosensor, can each be connected attheir second ends to a third fiber probe that can act as both anilluminator and a collector. The third fiber probe is oriented at anormal angle of incidence to the biosensor and supportscounter-propagating illuminating and reflecting optical signals. Anexample of such a detection system is shown in FIG. 18.

Another method of detection involves the use of a beam splitter thatenables a single illuminating fiber, which is connected to a lightsource, to be oriented at a 90 degree angle to a collecting fiber, whichis connected to a detector. Light is directed through the illuminatingfiber probe into the beam splitter, which directs light at thebiosensor. The reflected light is directed back into the beam splitter,which directs light into the collecting fiber probe. An example of sucha detection device is shown in FIG. 20. A beam splitter allows theilluminating light and the reflected light to share a common opticalpath between the beam splitter and the biosensor, so perfectlycollimated light can be used without defocusing.

Angular Scanning

Detection systems of the invention are based on collimated white lightillumination of a biosensor surface and optical spectroscopy measurementof the resonance peak of the reflected beam. Molecular binding on thesurface of a biosensor is indicated by a shift in the peak wavelengthvalue, while an increase in the wavelength corresponds to an increase inmolecular absorption.

As shown in theoretical modeling and experimental data, the resonancepeak wavelength is strongly dependent on the incident angle of thedetection light beam. FIG. 19 depicts this dependence as modeled for abiosensor of the invention. Because of the angular dependence of theresonance peak wavelength, the incident white light needs to be wellcollimated. Angular dispersion of the light beam broadens the resonancepeak, and reduces biosensor detection sensitivity. In addition, thesignal quality from the spectroscopic measurement depends on the powerof the light source and the sensitivity of the detector. In order toobtain a high signal-to-noise ratio, an excessively long integrationtime for each detection location can be required, thus lengtheningoverall time to readout a biosensor plate. A tunable laser source can beused for detection of grating resonance, but is expensive.

In one embodiment of the invention, these disadvantages are addressed byusing a laser beam for illumination of a biosensor, and a light detectorfor measurement of reflected beam power. A scanning mirror device can beused for varying the incident angle of the laser beam, and an opticalsystem is used for maintaining collimation of the incident laser beam.See, e.g., “Optical Scanning” (Gerald F. Marchall ed., Marcel Dekker(1991). Any type of laser scanning can be used. For example, a scanningdevice that can generate scan lines at a rate of about 2 lines to about1,000 lines per second is useful in the invention. In one embodiment ofthe invention, a scanning device scans from about 50 lines to about 300lines per second.

In one embodiment, the reflected light beam passes through part of thelaser scanning optical system, and is measured by a single lightdetector. The laser source can be a diode laser with a wavelength of,for example, 780 nm, 785 nm, 810 nm, or 830 nm. Laser diodes such asthese are readily available at power levels up to 150 mW, and theirwavelengths correspond to high sensitivity of Si photodiodes. Thedetector thus can be based on photodiode biosensors. An example of sucha detection system is shown in FIG. 52. A light source (100) provideslight to a scanner device (200), which directs the light into an opticalsystem (300) The optical system (300) directs light to a biosensor (400)Light is reflected from the biosensor (400) to the optical system (300),which then directs the light into a light signal detector (500). Oneembodiment of a detection system is shown in FIG. 21, which demonstratesthat while the scanning mirror changes its angular position, theincident angle of the laser beam on the surface changes by nominallytwice the mirror angular displacement. The scanning mirror device can bea linear galvanometer, operating at a frequency of about 2 Hz up toabout 120 Hz, and mechanical scan angle of about 10 degrees to about 20degrees. In this example, a single scan can be completed within about 10msec. A resonant galvanometer or a polygon scanner can also be used. Theexample shown in FIG. 21 includes a simple optical system for angularscanning. It consists of a pair of lenses with a common focal pointbetween them. The optical system can be designed to achieve optimizedperformance for laser collimation and collection of reflected lightbeam.

The angular resolution depends on the galvanometer specification, andreflected light sampling frequency. Assuming galvanometer resolution of30 arcsec mechanical, corresponding resolution for biosensor angularscan is 60 arcsec, i.e. 0.017 degree. In addition, assume a samplingrate of 100 ksamples/sec, and 20 degrees scan within 10 msec. As aresult, the quantization step is 20 degrees for 1000 samples, i.e. 0.02degree per sample. In this example, a resonance peak width of 0.2degree, as shown by Peng and Morris (Experimental demonstration ofresonant anomalies in diffraction from two-dimensional gratings, OpticsLett., 21:549 (1996)), will be covered by 10 data points, each of whichcorresponds to resolution of the detection system.

The advantages of such a detection system includes: excellentcollimation of incident light by a laser beam, high signal-to-noiseratio due to high beam power of a laser diode, low cost due to a singleelement light detector instead of a spectrometer, and high resolution ofresonance peak due to angular scanning.

Fiber Probe Biosensor

A biosensor of the invention can occur on the tip of a multi-mode fiberoptic probe. This fiber optic probe allows for in vivo detection ofbiomarkers for diseases and conditions, such as, for example, cardiacartery disease, cancer, inflammation, and sepsis. A single biosensorelement (comprising, for example, several hundred grating periods) canbe fabricated into the tip of a fiber optic probe, or fabricated from aglass substrate and attached to the tip of a fiber optic probe. See FIG.17. A single fiber is used to provide illumination and measure resonantreflected signal.

For example, a fiber probe structure similar to that shown in FIG. 18can be used to couple an illuminating fiber and detecting fiber into asingle counterpropagating fiber with a biosensor embedded or attached toits tip. The fiber optic probe is inserted into a mammalian body, forexample, a human body. Illumination and detection of a reflected signalcan occur while the probe is inserted in the body.

Mathematical Resonant Peak Determination

The sensitivity of a biosensor is determined by the shift in thelocation of the resonant peak when material is bound to the biosensorsurface. Because of noise inherent in the spectrum, it is preferable touse a procedure for determining an analytical curve—the turning point(i.e., peak) of which is well defined. Furthermore, the peakcorresponding to an analytic expression can be preferably determined togreater than sub-sampling-interval accuracy, providing even greatersensitivity.

One embodiment of the invention provides a method for determining alocation of a resonant peak for a binding partner in a resonantreflectance spectrum with a colormetric resonant biosensor. The methodcomprises selecting a set of resonant reflectance data for a pluralityof colormetric resonant biosensors or a plurality of biosensor distinctlocations. The set of resonant reflectance data is collected byilluminating a colormetric resonant diffractive grating surface with alight source and measuring reflected light at a pre-determinedincidence. The colormetric resonant diffractive grating surface is usedas a surface binding platform for one or more specific bindingsubstances such that binding partners can be detected without use of amolecular label.

The step of selecting a set of resonant reflectance data can includeselecting a set of resonant reflectance data:

x_(i) and y_(i) for i=1, 2, 3, . . . n,

wherein x_(i) is where a first measurement includes a first reflectancespectra of one or more specific binding substances attached to thecolormetric resonant diffractive grating surface, y_(i) and a secondmeasurement and includes a second reflectance spectra of the one or morespecific binding substances after a plurality of binding partners areapplied to colormetric resonant diffractive grating surface includingthe one or more specific binding substances, and n is a total number ofmeasurements collected.

The set of resonant reflectance data includes a plurality of sets of twomeasurements, where a first measurement includes a first reflectancespectra of one or more specific binding substances that are attached tothe colormetric resonant diffractive grating surface and a secondmeasurement includes a second reflectance spectra of the one or morespecific binding substances after one or more binding partners areapplied to the colormetric resonant diffractive grating surfaceincluding the one or more specific binding substances. A difference in apeak wavelength between the first and second measurement is ameasurement of an amount of binding partners that bound to the one ormore specific binding substances. A sensitivity of a colormetricresonant biosensor can be determined by a shift in a location of aresonant peak in the plurality of sets of two measurements in the set ofresonant reflectance data.

A maximum value for a second measurement from the plurality of sets oftwo measurements is determined from the set of resonant reflectance datafor the plurality of binding partners, wherein the maximum valueincludes inherent noise included in the resonant reflectance data. Amaximum value for a second measurement can include determining a maximumvalue y_(k) such that:

(y_(k)>=y_(i)) for all i≠k.

It is determined whether the maximum value is greater than apre-determined threshold. This can be calculated by, for example,computing a mean of the set of resonant reflectance data; computing astandard deviation of the set of resonant reflectance data; anddetermining whether ((y_(k)−mean)/standard deviation) is greater than apre-determined threshold. The pre-determined threshold is determined bythe user. The user will determine what amount of sensitivity is desiredand will set the pre-determined threshold accordingly.

If the maximum value is greater than a pre-determined threshold acurve-fit region around the determined maximum value is defined. Thestep of defining a curve-fit region around the determined maximum valuecan include, for example:

defining a curve-fit region of (2w+1) bins, wherein w is apre-determined accuracy value;

extracting (x_(i,), k−w<=i<=k+w); and

extracting (y_(i,), k−w<=i<=k+w).

A curve-fitting procedure is performed to fit a curve around thecurve-fit region, wherein the curve-fitting procedure removes apre-determined amount of inherent noise included in the resonantreflectance data. A curve-fitting procedure can include, for example:

computing g_(i)=ln y_(i);

performing a 2^(nd) order polynomial fit on g_(i) to obtain g′_(i)defined on

(x_(i,), k−w<=i<=k+w);

determining from the 2^(nd) order polynomial fit coefficients a, b and cof for (ax²+bx+c)−; and

computing y′_(i)=e^(g′i).

The location of a maximum resonant peak is determined on the fittedcurve, which can include, for example, determining a location of maximumreasonant peak (x_(p)=(−b)/2a). A value of the maximum resonant peak isdetermined, wherein the value of the maximum resonant peak is used toidentify an amount of biomolecular binding of the one or more specificbinding substances to the one or more binding partners. A value of themaximum resonant peak can include, for example, determining the valuewith of x_(p) at y′_(p).

One embodiment of the invention comprises a computer readable mediumhaving stored therein instructions for causing a processor to execute amethod for determining a location of a resonant peak for a bindingpartner in a resonant reflectance spectrum with a colormetric resonantbiosensor. A computer readable medium can include, for example, magneticdisks, optical disks, organic memory, and any other volatile (e.g.,Random Access Memory (“RAM”)) or non-volatile (e.g., Read-Only Memory(“ROM”)) mass storage system readable by the processor. The computerreadable medium includes cooperating or interconnected computer readablemedium, which exist exclusively on a processing system or to bedistributed among multiple interconnected processing systems that can belocal or remote to the processing system.

The following are provided for exemplification purpose only and are notintended to limit the scope of the invention described in broad termsabove. All references cited in this disclosure are incorporated hereinby reference.

EXAMPLE 1 Fabrication of a SWS Biosensor

An example of biosensor fabrication begins with a flat glass substratethat is coated with a thin layer (180 nm) of silicon nitride byplasma-enhanced chemical vapor deposition (PECVD).

The desired structure is first produced in photoresist by coherentlyexposing a thin photoresist film to three laser beams, as described inpreviously (Cowen, “The recording and large scale replication of crossedholographic grating arrays using multiple beam interferometry,” inInternational Conference on the Application, Theory, and Fabrication ofPeriodic Structures, Diffraction Gratings, and Moire Phenomena II, J. M.Lerner, ed., Proc. Soc. Photo-Opt. Instrum. Eng., 503, 120-129, 1984;Cowen, “Holographic honeycomb microlens,” Opt. Eng. 24, 796-802 (1985);Cowen & Slafer, “The recording and replication of holographicmicropatterns for the ordering of photographic emulsion grains in filmsystems,” J. Imaging Sci. 31, 100-107, 1987. The nonlinear etchingcharacteristics of photoresist are used to develop the exposed film tocreate a pattern of holes within a hexagonal grid, as shown in FIG. 22.The photoresist pattern is transferred into the silicon nitride layerusing reactive ion etching (RIE). The photoresist is removed, and acover layer of spin-on-glass (SOG) is applied (Honeywell ElectronicMaterials, Sunnyvale, Calif.) to fill in the open regions of the siliconnitride grating. The structure of the top surface of the finishedbiosensor is shown in FIG. 23. A photograph of finished parts are shownin FIG. 24.

EXAMPLE 2

A SRVD biosensor was prepared by making five circular diffuse gratingholograms by stamping a metal master plate into vinyl. The circularholograms were cut out and glued to glass slides. The slides were coatedwith 1000 angstroms of aluminum. In air, the resonant wavelength of thegrating is ˜380 nm, and therefore, no reflected color is visible. Whenthe grating is covered with water, a light blue reflection is observed.Reflected wavelength shifts are observable and measurable while thegrating is covered with a liquid, or if a specific binding substancesand/or binding partners cover the structure.

Both proteins and bacteria were immobilized onto the surface of a SRVDbiosensor at high concentration and the wavelength shift was measured.For each material, a 20 μl droplet is placed onto a biosensor distinctlocation and allowed to dry in air. At 1 μg/ml protein concentration, a20 μl droplet spreads out to cover a 1 cm diameter circle and depositsabout 2×10⁻⁸ grams of material. The surface density is 25.6 ng/mm².

For high concentration protein immobilization (biosensor 4) a 10 μldroplet of 0.8 g bovine serum albumin (BSA) in 40 ml DI H₂O is spreadout to cover a 1 cm diameter circle on the surface of a biosensor. Thedroplet deposits 0.0002 g of BSA, for a density of 2.5e-6 g/mm². Afterprotein deposition, biosensor 4 has a green resonance in air.

For bacteria immobilization (biosensor 2) a 20 μl droplet of NECKborrelia Lyme Disease bacteria (1.8e8 cfu/ml) was deposited on thesurface of a biosensor. After bacteria deposition, the biosensor looksgrey in air.

For low concentration protein immobilization (biosensor 6) a 10 μldroplet of 0.02% of BSA in DI H₂O (0.8 g BSA in 40 ml DI H₂O) is spreadout to cover a 1 cm diameter circle. The droplet deposits 0.000002 g ofBSA for a density of 2.5e-8 g/mm². After protein deposition, biosensor 6looks grey in air.

In order to obtain quantitative data on the extent of surfacemodification resulting from the above treatments, the biosensors weremeasured using a spectrometer.

Because a green resonance signal was immediately visually observed onthe biosensor upon which high concentration BSA was deposited (biosensor4), it was measured in air. FIG. 25 shows two peaks at 540 nm and 550 nmin green wavelengths where none were present before protein deposition,indicating that the presence of a protein thin film is sufficient toresult in a strong shift in resonant wavelength of a surface reliefstructure.

Because no visible resonant wavelength was observed in air for the slideupon which a low concentration of protein was applied (biosensor 6), itwas measured with distilled water on surface and compared against abiosensor which had no protein treatment. FIG. 26 shows that theresonant wavelength for the slide with protein applied shifted to greencompared to a water-coated slide that had not been treated.

Finally, a water droplet containing Lyme Disease bacteria Borreliaburgdorferi was applied to a grating structure and allowed to dry in air(biosensor 2). Because no visually observed resonance occurred in airafter bacteria deposition, the biosensor was measured with distilledwater on the surface and compared to a water-coated biosensor that hadundergone no other treatment. As shown in FIG. 27, the application ofbacteria results in a resonant frequency shift to longer wavelengths.

EXAMPLE 3 Computer Model of Biosensor

To demonstrate the concept that a resonant grating structure can be usedas a biosensor by measuring the reflected wavelength shift that isinduced when biological material is adsorbed onto its surface, thestructure shown in FIG. 1 was modeled by computer. For purposes ofdemonstration, the substrate chosen was glass (n_(substrate)=1.50). Thegrating is a two-dimensional pattern of silicon nitride squares (t₂=180nm, n₂=2.01, k₂=0.001) with a period of 510 nm, and a filling factor of56.2% (i.e. 56.2% of the surface is covered with silicon nitride squareswhile the rest is the area between the squares). The areas betweensilicon nitride squares are filled with a lower refractive indexmaterial. The same material also covers the squares and provides auniformly flat upper surface. For this simulation, a glass layer wasselected (n₁=1.40) that covers the silicon nitride squares by t₂=100 nm.To observe the effect on the reflected wavelength of this structure withthe deposition of biological material, variable thicknesses of protein(n_(bio)=1.5) were added above the glass coating layer.

The reflected intensity as a function of wavelength was modeled usingGSOLVER software, which utilizes full 3-dimensional vector code usinghybrid Rigorous Coupled Wave Analysis and Modal analysis. GSOLVERcalculates diffracted fields and diffraction efficiencies from planewave illumination of arbitrarily complex grating structures. Theillumination may be from any incidence and any polarization.

The results of the computer simulation are shown in FIG. 28 and FIG. 29.As shown in FIG. 28, the resonant structure allows only a singlewavelength, near 780 nm, to be reflected from the surface when noprotein is present on the surface. Because the peak width athalf-maximum is ˜1.5 nm, resonant wavelength shifts of 0.2 nm will beeasily resolved. FIG. 28 also shows that the resonant wavelength shiftsto longer wavelengths as more protein is deposited on the surface of thestructure. Protein thickness changes of 2 nm are easily observed. FIG.29 plots the dependence of resonant wavelength on the protein coatingthickness. A near linear relationship between protein thickness andresonant wavelength is observed, indicating that this method ofmeasuring protein adsorption can provide quantitative data. For thesimulated structure, FIG. 29 shows that the wavelength shift responsebecomes saturated when the total deposited protein layer exceeds 250 nm.This upper limit for detection of deposited material provides adequatedynamic range for any type of biomolecular assay.

EXAMPLE 4 Computer Model of Biosensor

In another embodiment of the invention a biosensor structure shown inFIG. 30 was modeled by computer. For purposes of demonstration, thesubstrate chosen was glass n_(substrate)=1.454 coated with a layer ofhigh refractive index material such as silicon nitride, zinc sulfide,tantalum oxide, or titanium dioxide. In this case, silicon nitride(t₃=90 nm, n₃=2.02) was used. The grating is two-dimensional pattern ofphotoresist squares (t₂=90 nm, n₂=1.625) with a period of 510 nm, and afilling factor of 56.2% (i.e. 56.2% of the surface is covered withphotoresist squares while the rest is the area between the squares). Theareas between photoresist squares are filled with a lower refractiveindex material such as glass, plastic, or epoxy. The same material alsocovers the squares and provides a uniformly flat upper surface. For thissimulation, a glass layer was selected (n₁=1.45) that covers thephotoresist squares by t₂=100 nm. To observe the effect on the reflectedwavelength of this structure with the deposition of a specific bindingsubstance, variable thicknesses of protein (n_(bio)=1.5) were addedabove the glass coating layer.

The reflected intensity as a function of wavelength was modeled usingGSOLVER software, which utilizes full 3-dimensional vector code usinghybrid Rigorous Coupled Wave Analysis and Modal analysis. GSOLVERcalculates diffracted fields and diffraction efficiencies from planewave illumination of arbitrarily complex grating structures. Theillumination may be from any incidence and any polarization.

The results of the computer simulation are shown in FIG. 31 and FIG. 32.The resonant structure allows only a single wavelength, near 805 nm, tobe reflected from the surface when no protein is present on the surface.Because the peak width at half-maximum is <0.25 nm, resonant wavelengthshifts of 1.0 nm will be easily resolved. FIG. 31 also shows that theresonant wavelength shifts to longer wavelengths as more protein isdeposited on the surface of the structure. Protein thickness changes of1 nm are easily observed. FIG. 32 plots the dependence of resonantwavelength on the protein coating thickness. A near linear relationshipbetween protein thickness and resonant wavelength is observed,indicating that this method of measuring protein adsorption can providequantitative data.

EXAMPLE 5 Sensor Readout Instrumentation

In order to detect reflected resonance, a white light source canilluminate a 1 mm diameter region of a biosensor surface through a 400micrometer diameter fiber optic and a collimating lens, as shown in FIG.33. Smaller or larger areas may be sampled through the use ofillumination apertures and different lenses. A group of six detectionfibers are bundled around the illumination fiber for gathering reflectedlight for analysis with a spectrometer (Ocean Optics, Dunedin, Fla.).For example, a spectrometer can be centered at a wavelength of 800 nm,with a resolution of ˜0.14 nm between sampling bins. The spectrometerintegrates reflected signal for 25-75 msec for each measurement. Thebiosensor sits upon an x-y motion stage so that different regions of thebiosensor surface can be addressed in sequence.

Equivalent measurements can be made by either illuminating the topsurface of device, or by illuminating through the bottom surface of thetransparent substrate. Illumination through the back is preferred whenthe biosensor surface is immersed in liquid, and is most compatible withmeasurement of the biosensor when it is incorporated into the bottomsurface of, for example, a microwell plate.

EXAMPLE 6 Demonstration of Resonant Reflection

FIG. 34 shows the resonant reflectance spectra taken from a biosensor asshown in FIG. 1 using the instrumentation described in Example 5. Thewavelength of the resonance (λ_(peak)=772.5 nm) compares with theresonant wavelength predicted by the computer model (λ_(peak)=781 nm),and the measured reflectance efficiency (51%) is comparable to thepredicted efficiency (70%). The greatest discrepancy between themeasured and predicted characteristics is the linewidth of the resonantpeak. The measured full-width at half maximum (FWHM) of the resonance is6 nm, while the predicted FWHM is 1.5 nm. As will be shown, the dominantsource of the larger measured FWHM is collimation of the illuminationoptics, which can easily be corrected.

As a basic demonstration of the resonant structure's ability to detectdifferences in the refractive index of materials in contact with itssurface, a biosensor was exposed to a series of liquids withwell-characterized optical properties. The liquids used were water,methanol, isopropyl alcohol, acetone, and DMF. A biosensor was placedface-down in a small droplet of each liquid, and the resonant wavelengthwas measured with a fiber illumination/detection probe facing thebiosensor's back side. Table 1 shows the calculated and measured peakresonant wavelength as a biosensor surface is exposed to liquids withvariable refractive index demonstrating the correlation between measuredand theoretical detection sensitivity. As shown in Table 1, the measuredresonant peak positions and measured resonant wavelength shifts arenearly identical to the predicted values. This example demonstrates theunderlying sensitivity of the biosensor, and validates the computermodel that predicts the wavelength shift due to changes in the materialin contact with the surface.

TABLE 1 Calculated Measured Peak Peak Wavelength Wavelength Solution n(nm) Shift (nm) (nm) Shift (nm) Water 1.333 791.6 0 786.08 0 Isopropyl1.3776 795.9 4.3 789.35 3.27 Acetone 1.3588 794 2.4 788.22 2.14 Methanol1.3288 791.2 −0.4 785.23 −0.85 DMF 1.4305 802 10.4 796.41 10.33

Similarly, a biosensor is able to measure the refractive indexdifference between various buffer solutions. As an example, FIG. 35shows the variation in peak wavelength with the concentration of bovineserum albumin (BSA) in water. Resonance was measured with the biosensorplaced face-down in a droplet of buffer, and rinsed with water betweeneach measurement.

EXAMPLE 7 Immobilized Protein Detection

While the detection experiments shown in Example 6 demonstrate abiosensor's ability to measure small differences in refractive index ofliquid solutions, the biosensor is intended to measure specific bindingsubstances and binding partners that are chemically bound to thebiosensor surface. In order to demonstrate a biosensor's ability toquantify biomolecules on its surface, droplets of BSA dissolved in PBSat various concentrations were applied to a biosensor as shown inFIG. 1. The 3 μl droplets were allowed to dry in air, leaving a smallquantity of BSA distributed over a ˜2 mm diameter area. The peakresonant wavelength of each biosensor location was measured before andafter droplet deposition, and the peak wavelength shift was recorded.See FIG. 37.

EXAMPLE 8 Immobilization of One or More Specific Binding Substances

The following protocol was used on a calorimetric resonant reflectivebiosensor to activate the surface with amine functional groups. Aminegroups can be used as a general-purpose surface for subsequent covalentbinding of several types of linker molecules.

A biosensor of the invention is cleaned by immersing it into piranhaetch (70/30% (v/v) concentrated sulfuric acid/30% hydrogen peroxide) for12 hours. The biosensor was washed thoroughly with water. The biosensorwas dipped in 3% 3-aminopropyltriethoxysilane solution in dry acetonefor 1 minute and then rinsed with dry acetone and air-dried. Thebiosensor was then washed with water.

A semi-quantitative method is used to verify the presence of aminogroups on the biosensor surface. One biosensor from each batch ofamino-functionalized biosensors is washed briefly with 5 mL of 50 mMsodium bicarbonate, pH 8.5. The biosensor is then dipped in 5 mL of 50mM sodium bicarbonate, pH 8.5 containing 0.1 mMsulfo-succinimidyl-4-O-(4,4′-dimethoxytrityl)-butyrate (s-SDTB, Pierce,Rockford, Ill.) and shaken vigorously for 30 minutes. The s-SDTBsolution is prepared by dissolving 3.0 mg of s-SDTB in 1 mL of DMF anddiluting to 50 mL with 50 mM sodium bicarbonate, pH 8.5. After a 30minute incubation, the biosensor is washed three times with 20 mL ofddH₂O and subsequently treated with 5 mL 30% perchloric acid. Thedevelopment of an orange-colored solution indicates that the biosensorhas been successfully derivatized with amines; no color change isobserved for untreated glass biosensors.

The absorbance at 495 nm of the solution after perchloric acid treatmentfollowing the above procedure can be used as an indicator of thequantity of amine groups on the surface. In one set of experiment, theabsorbance was 0.627, 0.647, and 0.728 for Sigma slides, Cel-Associateslides, and in-house biosensor slides, respectively. This indicates thatthe level of NH₂ activation of the biosensor surface is comparable inthe activation commercially available microarray glass slides.

After following the above protocol for activating the biosensor withamine, a linker molecule can be attached to the biosensor. Whenselecting a cross-linking reagent, issues such as selectivity of thereactive groups, spacer arm length, solubility, and cleavability shouldbe considered. The linker molecule, in turn, binds the specific bindingsubstance that is used for specific recognition of a binding partner. Asan example, the protocol below has been used to bind a biotin linkermolecule to the amine-activated biosensor.

Protocol for Activating Amine-Coated Biosensor with Biotin

Wash an amine-coated biosensor with PBS (pH 8.0) three times. Preparesulfo-succinimidyl-6-(biotinamido)hexanoate (sulfo-NHS-LC-biotin,Pierce, Rockford, Ill.) solution in PBS buffer (pH 8) at 0.5 mg/mlconcentration. Add 2 ml of the sulfo-NHS-LC-biotin solution to eachamine-coated biosensor and incubate at room temperature for 30 min. Washthe biosensor three times with PBS (pH 8.0). The sulfo-NHS-LC-biotinlinker has a molecular weight of 556.58 and a length of 22.4 Å. Theresulting biosensors can be used for capturing avidin or strepavidinmolecules.

Protocol for Activating Amine-Coated Biosensor with Aldehyde

Prepare 2.5% glutaraldehyde solution in 0.1 M sodium phosphate, 0.05%sodium azide, 0.1% sodium cyanoborohydride, pH 7.0. Add 2 ml of theglutaraldehyde solution to each amine-coated biosensor and incubate atroom temperature for 30 min. Wash the biosensor three times with PBS (pH7.0). The glutaraldehyde linker has a molecular weight of 100.11. Theresulting biosensors can be used for binding proteins and otheramine-containing molecules. The reaction proceeds through the formationof Schiff bases, and subsequent reductive amination yields stablesecondary amine linkages. In one experiment, where a coated aldehydeslide made by the inventors was compared to a commercially availablealdehyde slide (Cel-Associate), ten times higher binding of streptavidinand anti-rabbit IgG on the slide made by the inventors was observed.

Protocol for Activating Amine-Coated Biosensor with NHS

25 mM N,N′-disuccinimidyl carbonate (DSC, Sigma Chemical Company, St.Louis, Mo.) in sodium carbonate buffer (pH 8.5) was prepared. 2 ml ofthe DSC solution was added to each amine-coated biosensor and incubatedat room temperature for 2 hours. The biosensors were washed three timeswith PBS (pH 8.5). A DSC linker has a molecular weight of 256.17.Resulting biosensors are used for binding to hydroxyl- oramine-containing molecules. This linker is one of the smallesthomobifunctional NHS ester cross-linking reagents available.

In addition to the protocols defined above, many additional surfaceactivation and molecular linker techniques have been reported thatoptimize assay performance for different types of biomolecules. Mostcommon of these are amine surfaces, aldehyde surfaces, and nickelsurfaces. The activated surfaces, in turn, can be used to attach severaldifferent types of chemical linkers to the biosensor surface, as shownin Table 2. While the amine surface is used to attach several types oflinker molecules, the aldehyde surface is used to bind proteinsdirectly, without an additional linker. A nickel surface is usedexclusively to bind molecules that have an incorporated histidine(“his”) tag. Detection of “his-tagged” molecules with a Nickel activatedsurface is well known (Sigal et al., Anal Chem. 68, 490 (1996)).

Table 2 demonstrates an example of the sequence of steps that are usedto prepare and use a biosensor, and various options that are availablefor surface activation chemistry, chemical linker molecules, specificbinding substances and binding partners molecules. Opportunities alsoexist for enhancing detected signal through amplification with largermolecules such as HRP or streptavidin and the use of polymer materialssuch as dextran or TSPS to increase surface area available for molecularbinding.

TABLE 2 Label Bare Surface Linker Receptor Detected Molecule SensorActivation Molecule Molecule Material (Optional) Glass Amino SMPT Smm'cules Peptide Enhance Polymers Aldehyde NHS-Biotin Peptide Med Proteinsensitivity optional to Ni DMP Med Protein Lrg Protein 1000x enhanceNNDC Lrg Protein • IgG HRP sensitivity His-tag • IgG Phage Streptavidin2-5x Others . . . cDNA Cell Dextran TSPS cDNA

EXAMPLE 9 IgG Assay

As an initial demonstration for detection of biochemical binding, anassay was performed in which a biosensor was prepared by activation withthe amino surface chemistry described in Example 8 followed byattachment of a biotin linker molecule. The biotin linker is used tocovalently bond a streptavidin receptor molecule to the surface byexposure to a 50 μg/ml concentration solution of streptavidin in PBS atroom temperature for 2-4 hours. The streptavidin receptor is capable ofbinding any biotinylated protein to the biosensor surface. For thisexample, 3 μl droplets of biotinylated anti-human IgG in phosphatebuffer solution (PBS) were deposited onto 4 separate locations on thebiosensor surface at a concentration of 200 μg/ml. The solution wasallowed to incubate on the biosensor for 60 min before rinsingthoroughly with PBS. The peak resonant wavelength of the 4 locationswere measured after biotin activation, after streptavidin receptorapplication, and after ah-IgG binding. FIG. 37 shows that the additionof streptavidin and ah-IgG both yield a clearly measurable increase inthe resonant wavelength.

EXAMPLE 10 Biotin/Streptavidin Assay

A series of assays were performed to detect streptavidin binding by abiotin receptor layer. A biosensor was first activated with aminochemistry, followed by attachment of a NHS-Biotin linker layer, aspreviously described. Next, 3 μl droplets of streptavidin in PBS wereapplied to the biosensor at various concentrations. The droplets wereallowed to incubate on the biosensor surface for 30 min beforethoroughly washing with PBS rinsing with DI water. The peak resonantwavelength was measured before and after streptavidin binding, and theresonant wavelength shifts are shown in FIG. 38. A linear relationshipbetween peak wavelength and streptavidin concentration was observed, andin this case the lowest streptavidin concentration measured was 0.2μg/ml. This concentration corresponds to a molarity of 3.3 nM.

EXAMPLE 11 Protein-Protein Binding Assay

An assay was performed to demonstrate detection of protein-proteininteractions. As described previously, a biosensor was activated withamino chemistry and an NHS-biotin linker layer. A goat anti-biotinantibody receptor layer was attached to the biotin linker by exposingthe biosensor to a 50 μg/ml concentration solution in PBS for 60 min atroom temperature followed by washing in PBS and rinsing with DI water.In order to prevent interaction of nonspecific proteins with unboundbiotin on the biosensor surface, the biosensor surface was exposed to a1% solution of bovine serum albumin (BSA) in PBS for 30 min. The intentof this step is to “block” unwanted proteins from interacting with thebiosensor. As shown in FIG. 39 a significant amount of BSA isincorporated into the receptor layer, as shown by the increase in peakwavelength that is induced. Following blocking, 3 μl droplets of variousconcentrations of anti-goat IgG were applied to separate locations onthe biosensor surface. The droplets were allowed to incubate for 30 minbefore thorough rinsing with DI water. The biosensor peak resonantwavelength was measured before blocking, after blocking, after receptorlayer binding, and after anti-goat IgG detection for each spot. FIG. 39shows that an anti-goat IgG concentration of 10 μg/ml yields an easilymeasurable wavelength shift.

EXAMPLE 12 Unlabeled ELISA Assay

Another application of a biosensor array platform is its ability toperform Enzyme-Linked Immunosorbent Assays (ELISA) without the need foran enzyme label, and subsequent interaction an enzyme-specific substrateto generate a colored dye. FIG. 40 shows the results of an experimentwhere a biosensor was prepared to detect interferon-γ (IFN-γ) with anIFN-γ antibody receptor molecule. The receptor molecule was covalentlyattached to an NH₂-activated biosensor surface with an SMPT linkermolecule (Pierce Chemical Company, Rockford, Ill.). The peak resonantwavelength shift for application of the NH₂, SMPT, and anti-human IFN-αreceptor molecules were measured for two adjacent locations on thebiosensor surface, as shown in FIG. 40. The two locations were exposedto two different protein solutions in PBS at a concentration of 100μg/ml. The first location was exposed to IFN-γ, which is expected tobind with the receptor molecule, while the second was exposed to neuralgrowth factor (NGF), which is not expected to bind with the receptor.Following a 30 minute incubation the biosensor was measured byilluminating from the bottom, while the top surface remained immersed inliquid. The location exposed to IFN-γ registered a wavelength shift of0.29 nm, while the location exposed to NGF registered a wavelength shiftof only 0.14 nm. Therefore, without the use of any type of enzyme labelor color-generating enzyme reaction, the biosensor was able todiscriminate between solutions containing different types of protein.

EXAMPLE 13 Protease Inhibitor Assay (Caspase-3)

A Caspase-3 protease inhibitor assay was performed to demonstrate thebiosensor's ability to measure the presence and cleavage of smallmolecules in an experimental context that is relevant to pharmaceuticalcompound screening.

Caspases (Cysteine-requiring Aspartate protease) are a family ofproteases that mediate cell death and are important in the process ofapoptosis. Caspase 3, an effector caspase, is the most studied ofmammalian caspases because it can specifically cleave most knowncaspase-related substrates. The caspase 3 assay is based on thehydrolysis of the 4-amino acid peptide substrate NHS-Gly-Asp-Glu-Val-Aspp-nitroanilide (NHS-GDEVD-pNA) by caspase 3, resulting in the release ofthe pNA moiety.

The NHS molecule attached to the N-terminal of the GDEVD provides areactive end group to enable the NHS-GDEVD-pNA complex to be covalentlybound to the biosensor with the pNA portion of the complex oriented awayfrom the surface. Attached in this way, the caspase-3 will have the bestaccess to its substrate cleavage site.

A biosensor was prepared by cleaning in 3:1 H₂SO₄:H₂O₂ solution (roomtemperature, 1 hour), followed by silanation (2% silane in dry acetone,30 sec) and attachment of a poly-phe-lysine (PPL) layer (100 μg/ml PPLin PBS pH 6.0 with 0.5 M NaCl, 10 hours). The NHS-GDEVD-pNA complex wasattached by exposing the biosensor to a 10 mM solution in PBS (pH 8.0,room temperature, 1 hour). A microwell chamber was sealed over thebiosensor surface, and cleavage of pNA was performed by addition of 100μl of caspase-3 in 1× enzyme buffer (100 ng/ml, room temperature, 90minutes). Following exposure to the caspase 3 solution, the biosensorwas washed in PBS. A separate set of experiments using aspectrophotometer were used to confirm the attachment of the complex tothe surface of the biosensor, and functional activity of the caspase-3for removal of the pNA molecule from the surface-bound complex.

The peak resonant frequency of the biosensor was measured beforeattachment of the NHS-GDEVD-pNA complex, after attachment of the complex(MW=860 Da), and after cleavage of the pNA (MW=136) with caspase 3. Asshown in FIG. 41, the attachment of the peptide molecule is clearlymeasurable, as is the subsequent removal of the pNA. The pNA removalsignal of Δλ=0.016 nm is 5.3× higher than the minimum detectable peakwavelength shift of 0.003 nm. The proportion of the added molecularweight and subtracted molecular weight (860 Da/136 Da=6.32) are in closeagreement with the proportion of peak wavelength shift observed for theadded and subtracted material (0.082 nm/0.016 nm=5.14).

The results of this experiment confirm that a biosensor is capable ofmeasuring small peptides (in this case, a 5-mer peptide) without labels,and even detecting the removal of 130 Da portions of a molecule throughthe activity of an enzyme.

EXAMPLE 14 Reaction Kinetics for Protein-Protein Binding Assays

Because a biosensor of the invention can be queried continuously as afunction of time while it is immersed in liquid, a biosensor can beutilized to perform both endpoint-detection experiments and to obtainkinetic information about biochemical reactions. As an example, FIG. 42shows the results of an experiment in which a single biosensor locationis measured continuously through the course of consecutively addingvarious binding partners to the surface. Throughout the experiment, adetection probe illuminated the biosensor through the back of thebiosensor substrate, while biochemistry is performed on the top surfaceof the device. A rubber gasket was sealed around the measured biosensorlocation so that added reagents would be confined, and all measurementswere performed while the top surface of the biosensor was immersed inbuffer solution. After initial cleaning, the biosensor was activatedwith NH₂, and an NHS-Biotin linker molecule. As shown in FIG. 42, goatα-biotin antibodies of several different concentrations (1, 10, 100,1000 μg/ml) were consecutively added to the biosensor and allowed toincubate for 30 minutes while the peak resonant wavelength wasmonitored. Following application of the highest concentration α-BiotinIgG, a second layer of protein was bound to the biosensor surfacethrough the addition of α-goat IgG at several concentrations (0.1, 1,10, and 100 μg/ml). Again, the resonant peak was continuously monitoredas each solution was allowed to incubate on the biosensor for 30minutes. FIG. 42 shows how the resonant peak shifted to greaterwavelength at the end of each incubation period.

FIG. 43 shows the kinetic binding curve for the final resonant peaktransitions from FIG. 42, in which 100 μg/ml of α-goat IgG is added tothe biosensor. The curve displays the type of profile that is typicallyobserved for kinetic binding experiments, in which a rapid increase fromthe base frequency is initially observed, followed by a gradualsaturation of the response. This type of reaction profile was observedfor all the transitions measured in the experiment. FIG. 44 shows thekinetic binding measurement of IgG binding.

The removal of material from the biosensor surface through the activityof an enzyme is also easily observed. When the biosensor from the aboveexperiment (with two protein coatings of goat anti-biotin IgG andanti-goat IgG) is exposed to the protease pepsin at a concentration of 1mg/ml, the enzyme dissociates both IgG molecules, and removes them fromthe biosensor surface. As shown in FIG. 45, the removal of boundmolecules from the surface can be observed as a function of time.

EXAMPLE 15 Proteomics Applications

Biosensors of the invention can be used for proteomics applications. Abiosensor array can be exposed to a test sample that contains a mixtureof binding partners comprising, for example, proteins or a phage displaylibrary, and then the biosensor surface is rinsed to remove all unboundmaterial. The biosensor is optically probed to determine which distinctlocations on the biosensor surface have experienced the greatest degreeof binding, and to provide a quantitative measure of bound material.Next, the biosensor is placed in a “flow cell” that allows a small(e.g., <50 microliters) fixed volume of fluid to make contact to thebiosensor surface. One electrode is activated so as to elute boundmaterial from only a selected biosensor array distinct location. Thebound material becomes diluted within the flow cell liquid. The flowcell liquid is pumped away from the biosensor surface and is storedwithin a microtiter plate or some other container. The flow cell liquidis replaced with fresh solution, and a new biosensor electrode isactivated to elute its bound binding partners. The process is repeateduntil all biosensor distinct locations of interest have been eluted andgathered into separate containers. If the test sample liquid contained amixture of proteins, protein contents within the separate containers canbe analyzed using a technique such as electrospray tandem massspectrometry. If the sample liquid contained a phage display library,the phage clones within the separate containers can be identifiedthrough incubation with a host strain bacteria, concentrationamplification, and analysis of the relevant library DNA sequence.

EXAMPLE 16 Mathematical Resonant Peak Determination

This example discusses some of the findings that have been obtained fromlooking at fitting different types of curves to the observed data.

The first analytic curve examined is a second-order polynomial, given by

y=ax ² +bx+c

The least-squares solution to this equation is given by the costfunction

${\varphi = {\sum\limits_{i = 1}^{n}\left( {{ax}_{i}^{2} + {bx}_{i} + c - y_{i}} \right)^{2}}},$

the minimization of which is imposed by the constraints

$\frac{\partial\varphi}{\partial a} = {\frac{\partial\varphi}{\partial b} = {\frac{\partial\varphi}{\partial c} = 0.}}$

Solving these constraints for a, b, and c yields

$\begin{pmatrix}a \\b \\c\end{pmatrix} = {\begin{pmatrix}{\sum x_{i}^{4}} & {\sum x_{i}^{3}} & {\sum x_{i}^{2}} \\{\sum x_{i}^{3}} & {\sum x_{i}^{2}} & {\sum x_{i}} \\{\sum{x\; 2}} & {\sum x_{i}} & n\end{pmatrix}^{- 1} \cdot {\begin{pmatrix}{\sum{x_{i}^{2}y_{i}}} \\{\sum{x_{i}y_{i}}} \\y_{i}\end{pmatrix}.}}$

The result of one such fit is shown in FIG. 46; the acquired data areshown as dots and the 2^(nd)-order polynomial curve fit is shown as thesolid line.

Empirically, the fitted curve does not appear to have sufficient riseand fall near the peak. An analytic curve that provides bettercharacteristics in this regard is the exponential, such as a Gaussian. Asimple method for performing a Gaussian-like fit is to assume that theform of the curve is given by

y=e ^(ax) ² ^(+bx+c),

in which case the quadratic equations above can be utilized by formingy′, where y′=lny. FIG. 46 shows the result of such a fit. The visualappearance of FIG. 46 indicates that the exponential is a better fit,providing a 20% improvement over that of the quadratic fit.

Assuming that the exponential curve is the preferred data fittingmethod, the robustness of the curve fit is examined in two ways: withrespect to shifts in the wavelength and with respect to errors in thesignal amplitude.

To examine the sensitivity of the analytical peak location when thewindow from which the curve fitting is performed is altered to fall 10sampling intervals to the left or to the right of the true maxima. Theresulting shift in mathematically-determined peak location is shown inTable 3. The conclusion to be derived is that the peak location isreasonably robust with respect to the particular window chosen: for ashift of ˜1.5 nm, the corresponding peak location changed by only <0.06nm, or 4 parts in one hundred sensitivity.

To examine the sensitivity of the peak location with respect to noise inthe data, a signal free of noise must be defined, and then incrementalamounts of noise is added to the signal and the impact of this noise onthe peak location is examined. The ideal signal, for purposes of thisexperiment, is the average of 10 resonant spectra acquisitions.

Gaussian noise of varying degrees is superimposed on the ideal signal.For each such manufactured noisy signal, the peak location is estimatedusing the 2^(nd)-order exponential curve fit. This is repeated 25 times,so that the average, maximum, and minimum peak locations are tabulated.This is repeated for a wide range of noise variances—from a variance of0 to a variance of 750. The result is shown in FIG. 47.

TABLE 3 Comparison of peak location as a function of window locationShift Window Peak Location Δ = −10 bins 771.25-782.79 nm 778.8221 nm Δ =0 bins 772.70-784.23 nm 778.8887 nm Δ = +10 bins 774.15-785.65 nm7778.9653 nm 

The conclusion of this experiment is that the peak location estimationroutine is extremely robust to noisy signals. The entire range of peaklocations in FIG. 47 is only 1.5 nm, even with as much random noisevariance of 750 superimposed—an amount of noise that is substantiallygreater that what has been observed on the biosensor thus far. Theaverage peak location, despite the level of noise, is within 0.1 nm ofthe ideal location.

Based on these results, a basic algorithm for mathematically determiningthe peak location of a calorimetric resonant biosensor is as follows:

1. Input data x_(i) and y_(i), i=1, . . . , n

2. Find maximum

-   -   a. Find k such that y_(k)≧y_(i) for all i≠k

3. Check that maximum is sufficiently high

-   -   a. Compute mean y and standard deviation σ of sample    -   b. Continue only if (y_(k)− y)/σ>UserThreshold

4. Define curve-fit region of 2w+1 bins (w defined by the user)

-   -   a. Extract x_(i),k−w≦i≦k+w    -   b. Extract y,k−w≦i≦k+w

5. Curve fit

-   -   a. g_(i)=ln y_(i)    -   b. Perform 2^(nd)-order polynomial fit to obtain g_(i)′ defined        on x_(i),k−w≦i≦k+w    -   c. Polynomial fit returns coefficients a,b,c of form ax²+bx+c    -   d. Exponentiate: y_(i)′=e^(g) ^(i) ^(′)

6. Output

-   -   a. Peak location p given by x_(p)=−b/2a    -   b. Peak value given by y_(p)′(x_(p))

In summary, a robust peak determination routine has been demonstrated;the statistical results indicate significant insensitivity to the noisein the signal, as well as to the windowing procedure that is used. Theseresults lead to the conclusion that, with reasonable noise statistics,that the peak location can be consistently determined in a majority ofcases to within a fraction of a nm, perhaps as low as 0.1 to 0.05 nm.

EXAMPLE 17 Homogenous Assay Demonstration

An SWS biosensor detects optical density of homogenous fluids that arein contact with its surface, and is able to differentiate fluids withrefractive indices that differ by as little as Δn=4×10⁻⁵. Because asolution containing two free non-interacting proteins has a refractiveindex that is different from a solution containing two bound interactingproteins, an SWS biosensor can measure when a protein-proteininteraction has occurred in solution without any kind of particle tag orchemical label.

Three test solutions were prepared for comparison:

1. Avidin in Phosphate Buffer Solution (PBS), (10 μg/ml)

2. Avidin (10 μg/ml)+Bovine Serum Albumin (BSA) (10 μg/ml) in PBS

3. Avidin (10 μg/ml)+Biotinylated BSA (b-BSA) (10 μg/ml) in PBS

A single SWS sensor was used for all measurements to eliminate anypossibility of cross-sensor bias. A 200 μl sample of each test solutionwas applied to the biosensor and allowed to equilibrate for 10 minutesbefore measurement of the SWS biosensor peak resonant wavelength value.Between samples, the biosensor was thoroughly washed with PBS.

The peak resonant wavelength values for the test solutions are plottedin FIG. 51. The avidin solution was taken as the baseline reference forcomparison to the Avidin+BSA and Avidin+b-BSA solutions. Addition of BSAto avidin results in only a small resonant wavelength increase, as thetwo proteins are not expected to interact. However, because biotin andavidin bind strongly (Kd=10⁻¹⁵M), the avidin+b-BSA solution will containlarger bound protein complexes. The peak resonant wavelength value ofthe avidin+b-BSA solution thus provides a large shift compared toavidin+BSA.

The difference in molecular weight between BSA (MW=66 KDa) and b-BSA(MW=68 KDa) is extremely small. Therefore, the differences measuredbetween a solution containing non-interacting proteins (avidin+BSA) andinteracting proteins (avidin+b-BSA) are attributable only to differencesin binding interaction between the two molecules. The bound molecularcomplex results in a solution with a different optical refractive indexthan the solution without bound complex. The optical refractive indexchange is measured by the SWS biosensor.

EXAMPLE 18 Sensor Design and Fabrication

A one-dimensional linear grating surface biosensor structure requires agrating with a period lower than the wavelength of the resonantlyreflected light (R. Magnusson, and S. S. Wang, “New principle foroptical filters,” Appl. Phys. Lett., 61, No. 9, p. 1022, August, 1992,S. Peng and G. M. Morris, “Resonant scattering from two-dimensionalgratings,” J. Opt. Soc. Am. A, Vol. 13, No. 5, p. 993, May 1996). Asshown in FIG. 53, a one-dimensional linear grating surface structure wasfabricated from a low refractive index material that was overcoated witha thin film of higher refractive index material. The grating structurewas microreplicated within a layer of cured epoxy.

A one-dimensional linear grating surface structure was produced on thesurface of a plastic substrate material as follows. First, an 8-inchdiameter silicon “master” wafer was produced. The 550 nm period lineargrating structure was defined in photoresist using deep-UVphotolithography by stepping and repeating the exposure of a 9 mmdiameter circular grating reticle over the surface of aphotoresist-coated silicon wafer, as shown in FIG. 54. The exposurestep/repeat procedure produced patterns for two standard format 96-wellmicrotiter plates with 8 rows and 12 columns each. The exposedphotoresist was developed, and the grating structure was permanentlytransferred to the silicon wafer using a reactive ion etch with a depthof 200 nm. After etching, the photoresist was removed.

The grating structure was replicated onto a 0.005 inch thick sheet ofpolycarbonate by distributing a thin layer of epoxy between the siliconmaster wafer and a section of the polycarbonate sheet. The liquid epoxyconforms to the shape of the master grating, and was subsequently curedby exposure to ultraviolet light. The cured epoxy preferentially adheresto the polycarbonate sheet, and is peeled away from the silicon wafer.

Sensor fabrication was completed by sputter deposition of 120 nmtantalum oxide on the cured epoxy grating surface. Following tantalumoxide deposition, 3×5-inch microtiter plate sections were cut from thesensor sheet, activated with amine functional groups and attached to thebottoms of bottomless 96-well microtiter plates (Corning Costar®,Cambridge, Mass. and Greiner, Longwood, Fla.) with epoxy.

Surface Activation and Attachment of Receptor Molecule After highrefractive index material deposition, biosensors are activated withamine functional groups to enable various bifunctional linker moleculesto be attached to the surface in a known orientation. Amine activationwas performed by immersion of the sensor in 10%3-aminopropyltriethoxysilane (Pierce) solution in ethanol (Aldrich) for1 min, followed by a brief ethanol rinse. Activated sensors were thendried at 70° C. for 10 min. Other surface activation molecules couldinclude, for example, COOH, CHO, polymer, and Poly-phe-lysine.

A simple, colorimetric method using a modified protocol from Pierce wasused to determine the density of amine groups on the surface. Theamine-activated biosensor was immersed in 0.1 mM ofsulfo-succinimidyl-4-O-(4,4′-dimethoxytrityl)-butyrate (s-SDTB, Pierce),solution made in 50 mM sodium bicarbonate (pH 8.5), and shakenvigorously for 30 minutes. The biosensor was then washed with deionizedwater and subsequently treated with 30% perchloric acid (Sigma). Thesolution turned orange when the biosensor was amine-activated orremained colorless otherwise. This method indicated that the surfacedensity of the amine groups is 2×10¹⁴ groups/cm².

A one-dimensional linear grating resonant biosensor was used for thedetection of a well-characterized protein-protein binding interaction.The protein-protein system selected for this study was detection ofanti-biotin IgG antibody using biotin immobilized on the biosensorsurface as a receptor molecule. Therefore, a protocol for immobilizationof biotin on the biosensor surface was developed that utilizes abifunctional polyethyleneglycol-N-hydrosuccinimide (NHS-PEG) linkermolecule (Shearwater Polymers, Inc.) to act as an intermediary betweenthe amine surface group and the biotin. The NHS-PEG molecule is designedspecifically to enable NHS to preferentially bind to the amine-activatedsurface, leaving the PEG portion of the molecule oriented away from thesurface. The NHS-PEG linker molecule serves to separate the biotinmolecule from the biosensor surface by a short distance so it can retainits conformation, and thus its affinity for other molecules. The PEGalso serves to prevent nonspecific binding of proteins to the biosensor.

After attachment of amine-activated biosensor sheets into the bottom ofmicrotiter plates, individual microtiter wells were prepared with threedifferent surface functional groups in order to provide sufficientexperimental controls for the detection of anti-biotin IgG. First,amine-activated surfaces were studied without additional modification.The amine-activated surface is expected to bind proteinsnonspecifically, but not with high affinity. Second, microtiter wellswith the NHS-PEG bifunctional linker molecule were prepared. The NHS-PEGmolecule is expected to provide a surface that does not bind protein.Third, microtiter wells with an NHS-PEG-Biotin linker molecule wereprepared. The NHS-PEG-Biotin molecule is expected to bind strongly toanti-biotin IgG.

To activate an amine-coated sensor with biotin, 2 ml of NHS-PEG-Biotin(Shearwater) solution in TPBS (a reference buffer solution of 0.01%Tween™ 20 in phosphate buffer solution, pH 8) at 1.0 mg/ml concentrationwas added to the biosensor surface, and incubated at 37° C. for 1 hour.An identical procedure was used for attachment of the NHS-PEG(Shearwater) molecule without biotin. All purchased reagents were usedas packaged.

96-Well Plate Scanner Instrument

A schematic diagram of the system used to illuminate the biosensor andto detect the reflected signal is shown in FIG. 55. In order to detectthe reflected resonance, a white light source illuminates a ˜1 mmdiameter region of the grating surface through a 100 micrometer diameterfiber optic and a collimating lens at nominally normal incidence throughthe bottom of the microtiter plate. After passing through thecollimating lens, incident light passes through a linear polarizingfilter so that the linear grating is excited only with light that ispolarized either parallel or perpendicular to the grating lines.Reflected light passes through the polarizing filter again on its wayback to the detection probe. A detection fiber is bundled with theillumination fiber for gathering reflected light for analysis with aspectrometer (Ocean Optics). A series of 8 illumination/detection headsare arranged in a linear fashion, so that reflection spectra aregathered from all 8 wells in a microtiter plate column at once. Themicrotiter plate sits upon a motion stage so that each column can beaddressed in sequence.

Reflected Resonance Signal and Response Uniformity

With water in the microtiter plate well, the p-polarized reflectedresonance spectrum is shown in FIG. 56. The measured peak wavelengthvalue (PWV) is 857 nm, and the full-width at half-maximum (FWHM) of theresonant peak is 1.8 nm. Note that compared to structures producedusing, for example, two-dimensional grid holes on a hexagonal grid, onlya single resonant peak is measured, with a linewidth that is >3× lower.Most importantly, the single narrow resonant peak characteristic isuniformly obtained within every well of a 96-well microtiter platebiosensor.

The ability of a biosensor to measure shifts in optical density on itssurface can be calibrated by measuring the biosensor PWV when twosolutions with known refractive index values are added to the microtiterplate wells, and by calculating the PWV shift (ΔPWV) between the twosolutions. To measure biosensor response uniformity across a plate, thePWV of all 96 wells were measured in water (n=1.333) and subsequently inglycerol (n=1.472). The shift coefficient, σ=ΔPWV/Δn, is defined to be afigure of merit for comparing the response of biosensors with differentdesigns or fabrication processes. The average shift in PWV from water toglycerol across all wells was 15.57 nm, providing a shift coefficient ofσ=112 nm. The standard deviation of σ across 96 sensor wells was 1.07nm, indicating a very high degree of biosensor uniformity across a largesurface area.

Protein-Protein Binding Assay

A protein-antibody affinity assay was performed to demonstrate operationof the plastic biosensor. A matrix of three separate sensor surfacestates (NH₂, NHS-PEG, NHS-PEG-Biotin) were prepared and exposed to 7concentrations of goat anti-biotin IgG (Sigma). Each matrix location wasmeasured within a separate microtiter plate well, for a total of 21wells measured simultaneously. Because the NHS-PEG wells are notexpected to bind protein, they provide a reference for canceling commonmode effects such as the effect of the refractive index of the testsample and environmental temperature variation during the course of anassay. Data are reported here without the use of any mathematicalcorrection.

FIG. 57 plots the PWV shift—referenced to a sensor with no chemicalfunctional groups immobilized, recorded due to attachment of NH₂,NH₂+(NHS-PEG), and NH₂+(NHS-PEG-Biotin) molecules to the biosensorsurface. The error bars indicate the standard deviation of the recordedPWV shift over 7 microtiter plate wells. The data indicates that thebiosensor can differentiate between a clean surface, and one withimmobilized NH₂, as well as clearly detecting the addition of theNHS-PEG (MW=2000 Da) molecule. The difference between surfaceimmobilized NHS-PEG and NHS-PEG-Biotin (MW=3400 Da) is also measurable.

FIG. 58 shows the PWV shift response as a function of time for thebiosensor wells when exposed to various concentrations of anti-biotinIgG (0-80 μg/ml) and allowed to incubate for 20 minutes. The NHS-PEGsurface (FIG. 58B) provides the lowest response, while theamine-activated surface (FIG. 58A) demonstrates a low level ofnonspecific interaction with the anti-biotin IgG at high concentrations.The NHS-PEG-Biotin surface (FIG. 58C) clearly demonstrates strongspecific interaction with the anti-biotin IgG—providing strong PWVshifts in proportion to the concentration of exposed anti-biotin IgG.

The PWV shift magnitudes after 20 minutes from FIG. 58C are plotted as afunction of anti-biotin IgG concentration in FIG. 59. A roughly linearcorrelation between the IgG concentration and the measured PWV shift isobserved, and the lowest concentration IgG solution (1.25 μg/ml, 8.33nM) is clearly measurable over the negative control PSB solution.

EXAMPLE 19 Biosensor Comprising a Surface Modification Layer

A surface modification layer can be added to a one-dimensional gratingor two-dimensional grating biosensor of the invention. A surfacemodification layer is added to the top surface of the high refractiveindex material or cover layer of a biosensor and is useful forimmobilization of specific binding substances to the surface of abiosensor. A surface modification layer can comprise silicon oxide,silicon oxynitride, borosilicate glass, phosophosilicate glass, pyrex,any other glass (including BK7, SF11, LaSF9, Ultran, FK3, FK5) or anyother metal oxide. The thickness of the surface modification layer canbe about 5 nm to about 15 nm. In one embodiment of the invention thehigh refractive index material is tantalum oxide.

Silicon oxide was coated onto a biosensor of the invention by DCsputtering. Other possible methods of coating include evaporation, laserablation, chemical vapor deposition, and plasma-enhanced chemical vapordeposition. NH₂ was added to a biosensor with the surface modificationlayer and to a biosensor without a surface modification layer. A NH₂reacting fluorescent dye was used to visualize NH₂ attachment to thebiosensors. See FIG. 60. The higher fluorescence intensity from thesurface modified layer indicates much higher NH₂ density.

EXAMPLE 20 Biosensor Design with High Stability in Aqueous Solutions

A biosensor of the invention can be exposed to aqueous solutions duringuse. Stability in aqueous solutions can be added to a biosensor byadding a interfacial layer under the high refractive index materiallayer. For example, where a plastic grating surface, e.g., a lowrefractive index grating material, is coated with a high refractiveindex material, an interfacial layer can be added between the highrefractive index material and the low refractive index material.

For example, a biosensor was constructed by adding an adhesion enhancinginterfacial layer between a plastic grating surface and a highrefractive index optical material. Adhesion refers to the ability of athin film material to remain firmly attached to the material it isdeposited upon over a wide range of environmental conditions. Forexample, for a biosensor structure used in the invention, a highrefractive index thin film (such as silicon nitride, zinc sulfide,tantalum oxide, or titanium oxide) is deposited upon a grating surfacestructure formed from a cured epoxy material. Without an adhesionenhancing layer (called a “tie layer” or “interfacial layer”), the highrefractive index material can possibly delaminate from the gratingsurface structure under stringent experimental conditions, such as along-period exposure to a liquid. The tie layer material is selected tohave strong adhesion properties to both the underlying material and thehigh refractive index deposited material. Generally, the thickness ofthe tie layer is selected to be extremely thin, so as not to disrupt theoptical properties of the structure that it is embedded within. Tielayer thickness can range from about 1 nm to about 200 nm, for example.In this example, the tie layer was a silicon oxide layer ofapproximately 5 nm in thickness, the plastic material was Polyethylene(PET), and the high refractive index optical material was tantalumoxide. Alternative materials for the tie layer are silicon oxynitride,borosilicate glass, phosophosilicate glass, pyrex, any other glass(including BK7, SF11, LaSF9, Ultran, FK3, FK5) or any other metal oxide.Biosensor stability performance was improved in aqueous solutions by theaddition of silicon oxide interfacial layer. See FIG. 61. The improvedbiosensor stability performance allowed for the enhancement of signal tonoise ratio, which translates into more sensitive detection limits.

1. A detection system comprising: (a) a biosensor, wherein the biosensorcomprises: (i) a one-dimensional grating surface structure comprised ofa material having a low refractive index; (2) a high refractive indexmaterial layer that is applied on top of the low refractive indexone-dimensional grating layer; and (3) one or more specific bindingsubstances immobilized on a surface of the high refractive index layeropposite of the one-dimensional grating surface structure comprised of amaterial having a low refractive index; wherein the one or more specificbinding substances do not comprise detection labels; wherein, when thebiosensor is illuminated a resonant grating effect is produced on areflected radiation spectrum; and wherein the cross-sectional period ofthe one-dimensional grating is less than the wavelength of the resonantgrating effect; (b) a light source that directs light to the biosensor;and (c) a detector that detects light reflected or transmitted from thebiosensor.
 2. The detection system of claim 1, wherein a polarizingfilter is present between the light source and the biosensor.
 3. Thedetection system of claim 1, further comprising a fiber probe comprisingone or more illuminating optical fibers that are connected at a firstend to the light source, and one or more collecting optical fibersconnected at a first end to the detector, wherein the second ends of theilluminating and collecting fibers are arranged in line with acollimating lens that focuses light onto the biosensor.
 4. The detectionsystem of claim 3, wherein the illuminating fiber and the collectingfiber are the same fiber.
 5. The detection system of claim 1, whereinthe light source illuminates the biosensor from its top surface or fromits bottom surface.
 6. The detection system of claim 1, wherein the oneor more specific binding substances are arranged in an array of distinctlocations on the high refractive index material.